Molecular wire injection sensors

ABSTRACT

Disclosed is a sensor for sensing the presence of an analyte component without relying on redox mediators. This sensor includes (a) a plurality of conductive polymer strands each having at least a first end and a second end and each aligned in a substantially common orientation; (b) a plurality of molecular recognition headgroups having an affinity for the analyte component and being attached to the first ends of the conductive polymer strands; and (c) an electrode substrate attached to the conductive polymer strands at the second ends. The electrode substrate is capable of reporting to an electronic circuit reception of mobile charge carriers (electrons or holes) from the conductive polymer strands. The electrode substrate may be a photovoltaic diode.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation-in-part, claiming priority under 35U.S.C. § 120, from U.S. patent application Ser. No. 10/770,914, filedFeb. 2, 2004, naming inventor Randy E. Keen, and titled “MOLECULAR WIREINJECTION SENSORS,” which is in turn a continuation of U.S. patentapplication Ser. No. 09/960,165, filed Sep. 20, 2001 naming inventorRandy E. Keen, which is now U.S. Pat. No. 6,699,667, issued Mar. 2,2004, which is a continuation-in-part of U.S. patent application Ser.No. 09/365,109 filed on Jul. 30, 1999, which is now U.S. Pat. No.6,326,215 issued on Dec. 4, 2001, which is a divisional of U.S.application Ser. No. 08/856,822, filed May 14, 1997, now U.S. Pat. No.6,060,327, issued May 9, 2000. Each of these applications isincorporated herein by reference for all purposes.

BACKGROUND OF THE INVENTION

The present invention relates to biosensors and chemical sensors. Moreparticularly, it relates to sensors having a chemical or biochemicalspecies detection group connected to an electronic circuit byelectrically conducting polymer strands.

Biosensors employing enzymes have been applied to the detection ofnumerous analyte species concentrations including glucose, cholesterol,or both glucose and cholesterol concentrations in whole blood samples.Such sensors and associated instruments employ an enzyme capable ofcatalyzing a reaction at a rate representative of the selected compoundconcentration in an assay mixture.

There are three general detection approaches employing a glucose enzymeelectrode. The first and earliest measures oxygen consumption. Theoxygen-sensing probe is an electrolytic cell with a gold (or platinum)cathode separated from a tubular silver anode by an epoxy casting. Theanode is electrically connected to the cathode by electrolytic gel, andthe entire chemical system is isolated from the environment by a thingas-permeable membrane (often Teflon). A potential of approximately 0.8V(from solid-state power supply) is applied between the electrodes. Theoxygen in the sample diffuses through the membrane and is reduced at thecathode with the formation of the oxidation product, silver oxide, atthe silver anode. The resultant current is proportional to the amount ofoxygen reduced. The analyzer unit operates over the range from 0.2 to 50ppm of dissolved oxygen. Gases that reduce at 0.8V will interfere; theseinclude the halogens and SO₂. H₂S contaminates the electrodes.

A second approach detects H₂O₂ production but requires an appliedpotential of approximately 0.65V (from solid-state power supply) appliedbetween the electrodes, one of which is inside a permselective membrane.The H₂0₂ in the sample diffuses through the permselective membrane (ifone is present) and is oxidized at the anode. Many metal, metalcomplexes, nonmetal, organic and biochemical species that oxidize atapproximately 0.65V will interfere; such as ascorbic acid, amines,hydrazines, thiol compounds, catechols, hydroquinones, ferrocenes, andmetalloporphyrins. The inside permselective membrane is not alwayscapable of removing the complicated mix of possible interferences fromthe analyte matrix.

A third approach takes advantage of the fact that the enzymatic reactionrequires two steps. First, the enzyme glucose oxidase (GOD) (EC 1.1.3.4)is reduced by glucose, then the reduced enzyme is oxidized to itsinitial form by an electron acceptor, i.e., a mediator. In naturalsystems, the mediator is oxygen. In biosensors, another mediatorcompound may be employed to transfer electrons between the enzyme and aconductive surface of an electrode at a rate representative of theenzyme catalyzed reaction rate when an appropriate potential is appliedto the particular redox mediator in use. Such biosensors may employamperometric measurements to determine glucose concentration in a wholeblood sample. This involves an integrated sample measurement of the areaunder the ampere versus time curve, corresponding to the amount ofglucose in the sample.

The mechanism by which a common amperometric sensor works is depicted inFIG. 1. A sensor 2 employs glucose oxidase (GOD), for example, as amolecular recognition group. Glucose oxidase catalyzes the oxidation ofglucose to gluconolactone in analyte 4. This reaction involves theFAD/FADH₂ redox center of the enzyme. Sensor 2 includes a molecularrecognition group, region 6, attached to an electrode 8. When glucose inanalyte 4 contacts GOD-FAD (glucose oxidase including the FAD redoxcenter) in region 6, it is oxidized to gluconolactone. At the same time,the GOD-FAD is reduced to GOD-FADH₂. This involves two electrons and twohydrogen ions being transferred to the FAD. Normally, in the absence ofa sensor mediator, the GOD-FADH₂ is reoxidized by atmospheric oxygen toGOD-FAD to complete the catalytic reaction. In the presence of amediator, however, the GOD-FADH₂ is sometimes reoxidized by a mediator(Mox). In this case, the GOD-FADH₂ releases two hydrogen ions to analyte4 and two electrons to the mediator. The resulting reduced mediator(Mred) may then be reoxidized by electrode 8 at an appropriatepotential. The reoxidation of the mediator is accompanied by thetransfer of an electron or electrons to electrode 8. This is the currentthat is monitored to provide a concentration of glucose.

In theory, a mediator may be any small molecule inorganic,organometallic or organic compounds, which are reduced by the enzyme,and oxidized by an appropriate applied potential at the electrodesurface. The mediator should be designed to rapidly and efficientlytransfer electrons between the enzyme and the electrode. Otherwise,ambient oxygen would oxidize nearly all of the reduced GOD and thedesired signal would be very weak. The mediator should also transfer atotal charge proportional to the glucose or cholesterol concentration inthe sample. The current which results from the mediator oxidation isknown as the Cottrell current which, when integrated with respect totime, gives the number of coulombs associated with the sensor reaction.The total coulombs passed is proportional to the amount of analyte.

Unfortunately, mediators are commonly provided as mobile “reagents”which diffuse to the enzyme where they are oxidized or reduced(depending upon the reaction catalyzed by the enzyme). The oxidized orreduced mediator then diffuses to the electrode surface where it gainsor loses an electron. Unfortunately, such mechanism is dependent uponthe continuing presence of recycled mobile mediators. As such compoundscan leak from the electrode surfaces, there may be a gradual depletionin available mediator and a consequent reduction in sensor sensitivity.Examples of diffusing redox mediators include dyes (e.g., methyleneblue), ferrocene derivatives (Cass, AEG; Davis, G; Francis, G D; Hill,HAO; Aston, W J; Higgins, I J; Plotkin, E V; Scott, LDL; Turner, APF:Ferrocene-Mediated Enzyme Electrode for Amperometric Determination ofGlucose. Anal. Chem. 56:667-671, 1984), components of conducting organicmetals and quinones.

Also, available sensors applying the above amperometric approach to thedetection of glucose, cholesterol, lactate, H₂O₂, NAD(P)H, alcohol, anda variety of other compounds in whole blood samples, can have otherserious complicating problems. For example, the percentage of sensorsurface area covered by blood can vary; sometimes the blood sample doesnot cover the entire electrode. This may be caused by a poorly adherentenzyme (often applied by spraying) thus allowing leakage of blood orother analytes along the edges of the electrode. A related problemresults from hydration of the reaction area prior to test. This dilutesthe ligand (e.g., glucose) concentration and therefore gives a lowerreading than would be accurately given by an unhydrated surface.

Further, the partial pressure of molecular oxygen (O₂) may complicatethe interpretation of sensor data. Molecular oxygen is the naturalelectron acceptor mediator of the enzyme glucose oxidase (GOD).Following oxidation of D-(+)-glucose by GOD, reduced glucose oxidase(GOD_(red)) will transfer electrons to O₂ forming H₂O₂ in the absence ofother mediators. In amperometric glucose biosensors described above, theunwanted O₂ side reaction competes with synthetic chemical mediators forelectrons supplied by the GOD_(red) enzyme. Calibration of GOD-basedbiosensors at different altitudes (i.e., different partial pressures ofO₂) may be a problem if electron transfer rates of selected syntheticchemical mediators are not orders of magnitude faster than the O₂ sidereaction.

Humidity (i.e., H₂O) may be another potential problem if mass action ofH₂O and O₂ present drives the enzyme catalyzed oxidation product ofD-gluconolactone in reverse back to the reduced starting material,D-(+)-glucose. Catalase, a common contaminant of glucose oxidasepreparations, may be driven in reverse by mass action of excess H₂O andO₂ producing 2 moles of H₂O₂. H₂O₂ buildup combined withD-gluconolactone could drive the glucose oxidase reaction in reverse bymass action back to D-(+)-glucose.

Other problems associated with known amperometric sensors include, forexample, (1) difficulty in fitting the Cottrell current curve (i.e.,ampere-time graph), (2) sampling with enough frequency to accuratelyobtain the time integral of Cottrell current, (3) high applied potentialat the electrode causing indiscriminate oxidation or reduction ofinterfering substances, and (4) complicated electronic circuitsrequiring potentiostat and galvinostat instrumentation.

Some of the above drawbacks of the current amperometric biosensors havebeen noted and analyzed (see, Schuhmann, W: Chap. 9. Conducting PolymersAnd Their Application In Amperometric Biosensors. In: DiagnosticBiosensor Polymers. ACS Symposium Series 556. Usmani, A M; Akmal, N;eds. American Chemical Society; Washington, D.C.; 1994; pp. 110-123).First, due to the fact that the active site of redox enzymes is ingeneral deeply buried within the protein shell, direct electron transferbetween enzymes and electrode surfaces is rarely encountered. This isespecially true for enzymes which are integrated within non-conductingpolymer membranes in front of the electrode surface. Hence, electrontransfer is usually performed according to a ‘shuttle’ mechanisminvolving free-diffusing electron-transferring redox species for examplethe natural electron acceptor O₂ or artificial redox mediators likeferrocene derivatives (Cass, AEG; Davis, G; Francis, G D; Hill, HAO;Aston, W J; Higgins, I J; Plotkin, E V; Scott, LDL; Turner, APF:Ferrocene-Mediated Enzyme Electrode for Amperometric Determination ofGlucose. Anal. Chem. 56:667-671, 1984), osmium complexes (Heller, A:Electrical Wiring of Redox Enzymes. Acc. Chem. Res. 23(5):128-134,1990), or quinones. Due to the necessity for the redox mediators todiffuse freely between the active sites of the enzymes and the electrodesurface, these electrodes show a limited long-term stability as aconsequence of the unavoidable leaking of the mediator from the sensorsurface. Additionally in the case of the natural redox couple O₂/H₂O₂,the sensor signal is dependent on the O₂ partial pressure, and a highoperation potential has to be applied to the working electrode givingrise to possible interferences from cooxidizable compounds. The seconddrawback is related to the fabrication of these sensors. The physicalassembling of an enzyme membrane and an electrode is extremely difficultto automate and thus in principal incompatible with microelectronicfabrication techniques. Additionally, the miniaturization as well as theintegration of individual biosensors into a miniaturized sensor array isimpossible with techniques which are mainly based on the manualdeposition of a droplet of the membrane-forming mixture onto theelectrode surface.

Consequently, the next generation of amperometric enzyme electrodes hasto be based on immobilization techniques which are compatible withmicroelectronic mass-production processes and easy to miniaturize.Additionally, the integration of all necessary sensor components on thesurface of the electrode has to prevent the leaking of enzymes andmediators simultaneously improving the electron-transfer pathway fromthe active site of the enzyme to the electrode surface.

In addition to amperometric mechanisms, which rely on detecting currentgenerated from faradaic reactions, a potentiometric mechanism may beemployed to sense analyte concentration. Potentiometric techniquesmonitor potential changes between a working electrode and a referenceelectrode in response to charged ion species generated from enzymereactions on the working electrode. A very common potentiometric sensoris the pH sensor which registers changes in hydrogen ion concentrationin an analyte. A microelectronic potentiometric biosensor, the FieldEffect Transistor (FET) biosensor, has generated some interest. In thisdesign, a receptor or molecular recognition species is coated on atransistor gate. When a ligand binds with the receptor, the gateelectrode potential shifts, thereby controlling the current flowingthrough the FET. This current is detected by a circuit which converts itto an observed ligand concentration. Observed problems withpotentiometric systems include, for example, (1) slow response of theelectrode (i.e., seconds), (2) complicated electronic circuits for threeelectrode (i.e., working, counter, and reference electrode)electrochemical systems requiring potentiostat instrumentation, (3) lowsensitivity, and (4) limited dynamic range.

Recently, two groups (Heller et al. and Skotheim et al.) have exploredand developed redox polymers that can shuttle electrons from the enzymeto the electrode. The groups have “wired” the enzyme to the electrodewith a long redox polymer having a dense array of electron relays. Eachrelay is a redox site bound to the polymer backbone. Electrons movealong the polymer by hopping from one redox appendage to the next. Thepolymer penetrates and binds the enzymes, and is also bound to theelectrode.

Heller et al. have conducted work on Os-containing redox polymers. Theyhave synthesized a large number of such Os-containing polymers andevaluated their electrochemical characteristics (Gregg, B A; Heller, A:Redox Polymer Films Containing Enzymes. 1. A Redox-Conducting EpoxyCement: Synthesis, Characterization, and Electrocatalytic Oxidation ofHydroquinone. J. Phys. Chem. 95:5970-5975, 1991). Their most stable andreproducible redox polymer is a poly(4-vinyl pyridine) to whichOs(bpy)₂Cl₂ has been attached to ⅙th of the pendant pyridine groups. Theresultant redox polymer is water insoluble. To make it water soluble andbiologically compatible, Heller et al. have partially quaternized theremaining pyridine pendants with 2-bromoethyl amine. The redox polymeris water soluble and the newly introduced amine groups can react with awater soluble epoxy e.g., polyethylene glycol diglycidyl ether and GODto produce a cross-linked biosensor coating-film. Such coating-filmsproduced high current densities and a linear response to glucose up to600 mg/dL (U.S. Pat. No. 5,262,035 to Gregg et al.).

Heller describes the electrical wiring of redox enzymes for use asamperometric biosensors (Heller, A: Electrical Wiring of Redox Enzymes.Acc. Chem. Res. 23(5):128-134, 1990). The Heller approach is animprovement over amperometric enzyme electrodes based on diffusing redoxmediators, including dyes, ferrocene derivatives, components ofconducting organic metals, and quinones, all described above. In theHeller approach, redox centers of a redox polymer polycation (e.g.,2[Os-(2,2′-bipyridine)₂(poly(vinylpyridine))Cl]^(1+/2+)) areelectrostatically and covalently bound to the enzyme and relayselectrons to the electrode, on which a segment of the polycation isadsorbed. Binding of the redox polymer polycation to the electrode canbe electrostatic when the electrode has a negative surface charge.

Fluctuations in current with partial pressure of oxygen (e.g., oxygenconcentration in blood), depend on the ratio of the rate of directelectroxidation of the FADH₂ centers to their rate of oxidation bymolecular oxygen, and therefore on the rate of electron transfer to, andthe electrical resistance of, the three-dimensional wired-enzymestructure. At high osmium-complex concentrations, and in sufficientlythin layers, the competition is won by electron transfer to theelectrode via the osmium centers, and the electrodes are relativelyinsensitive to oxygen (Heller, A: Electrical Wiring of Redox Enzymes.Acc. Chem. Res. 23(5):128-134, 1990. Gregg, B A; Heller, A: Cross-LinkedRedox Gels Containing Glucose Oxidase for Amperometric BiosensorApplications. Anal. Chem. 62:258-263, 1990. Surridge, N A; Diebold, E R;Chang, J; Neudeck, G W: Chap 5. Electron-Transport Rates In An EnzymeElectrode For Glucose. In: Diagnostic Biosensor Polymers. ACS SymposiumSeries 556. Usmani, A M; Akmal, N; eds. American Chemical Society;Washington, D.C.; 1994; pp. 47-70).

Electrodes based on conducting polypyrroles with ferrocenes also havebeen reported (Hale, P D; Inagaki, T; Karan, H I; Okamoto, Y; Skotheim,T A: A New Class of Amperometric Biosensor Incorporating a PolymericElectron-Transfer Mediator. J. Am. Chem. Soc. 111(9):3482-3484, 1989).

Skotheim et al. have used flexible polymer chains to act as relays.Their polymers provide communication between GOD's redox centers andelectrode. No mediation was found when ferrocene was attached to anon-silicone backbone. Their ferrocene-modified siloxane polymers weresaid to be stable and non-diffusing (Boguslavsky, L I; Hale, P D;Skotheim, T A; Karan, H I; Lee, H S; Okamoto, Y: Novel Biosensors ForSpecific Neurotransmitters Based On Flavoenzymes And Flexible RedoxPolymers. Polym. Mater. Sci. Eng. 64:322-323, 1991).

Unfortunately, the redox polymer systems of Heller et al. and Skotheimet al. have a limited electron transfer rate based on electron hoppingbetween dense electron relay pendant groups. Further, their “wire” redoxcenters must be designed to undergo reaction at a potential close tothat of the enzyme catalyzed reaction. The closer the potential is tothe redox potential of the enzyme itself, the lesser the likelihood thata potentially interfering substrate will be spuriously oxidized.Unfortunately, to address this issue limits the range of polymer redoxcouple and molecular headgroup combinations.

A fundamental presupposition for the construction of reagentlessamperometric enzyme electrodes is the design of a suitableelectron-transfer pathway from the active site of the enzyme to theelectrode surface. According to Marcus theory (Marcus, R A; Sutin, N:Electron Transfers In Chemistry And Biology. Biochim. Biophys. Acta811:265-322, 1985) a redox mediator with a low reorganization energyafter the electron transfer has to be able to penetrate into the activesite of the enzyme to shorten the distance between the prosthetic group(e.g., FAD/FADH₂) and the mediator. Hence, the rate constant of theelectron-transfer reaction can be increased. After this ‘first’ electrontransfer the redox equivalents have to be transported to the electrodesurface via mechanism having a rate constant which is in the range ofthe turnover rate of the enzyme. In the shuttle mechanism mentionedabove (having mobile mediators), the electron transport involvesdiffusion of redox mediators. In the “wired” redox polymer sensorsdescribed above, electron transport involves hopping from one redoxcenter to the next on the polymer backbone.

In a recent study, Aizawa et al. discuss a reversible electron transferbetween the prosthetic group of pyrrolo quinoline quinone (PQQ) enzyme(fructose dehydrogenase) and an electrode through a molecular interface(Aizawa, M; Khan, G F; Kobatake, E; Haruyama, T; Ikariyama, Y: Chap. 26.Molecular Interfacing of Enzymes on the Electrode Surface. In:Interfacial Design and Chemical Sensing. ACS Symposium Series 561.Mallouk, T E; Harrison, D J; eds. American Chemical Society, Washington,D.C., 1994, pp. 305-313). The PQQ moieties of randomly oriented fructosedehydrogenase (FDH) which are very close to the transducer electrode caneasily transfer their electrons to the electrode (Shinohara, H; Khan, GF; Ikariyama, Y; Aizawa, M: Electrochemical Oxidation and Reduction ofPQQ Using a Conducting Polypyrrole-Coated Electrode. J. Electroanal.Chem. 304:75-84, 1991. Khan, G F; Shinohara, H; Ikariyama, y; Aizawa, M:Electrochemical Behaviour of Monolayer Quinoprotein Adsorbed on theElectrode Surface. J. Electroanal Chem. 315:263-273, 1991). However, theprosthetic groups of FDH located far from the electrode can not providetheir electrons, as the distance from the electrode exceeds the maximumelectron transfer distance (˜25 Å). Therefore, to make the FDH (EC1.1.99.11, MW: 141,000) on the electrode surface electrochemicallyactive, Aizawa et al. introduced an ultrathin conductive polypyrrole(PP) membrane as a molecular interface as “wiring” to assist theelectron transfer from PQQ to the electrode. Unfortunately, the wiringused by Aizawa is randomly oriented and does not necessarily presentenzyme at optimal position with respect to the analyte.

What is needed is an improved sensor design that rapidly transferselectrons from headgroup redox reactions to an electrode, does not relyon a redox relay such as freely diffusing mediators, and optimallyorients the headgroup with respect to the analyte.

A great number of approaches for microfabrication of chemical sensorsare currently under way, particularly in the areas of field effecttransistor (FET)-based chemical sensors, metal oxide gas sensors, andbiosensors. Since Janata et al. first reported micro-enzyme electrodesbased on FET (Caras, S; Janata, J: Field Effect Transistor Sensitive toPenicillin. Anal. Chem. 52:1935-1937, 1980), a number of groups havebeen employing microfabrication techniques (e.g., photolithography) suchas those employed in semiconductor device technology to fabricatemicro-enzyme electrodes. Despite enormous efforts of many groups, theFET-based micro-enzyme electrodes of practical use have not beenrealized yet, largely because of the problems associated withpotentiometric methods general lack of a fast response, highsensitivity, and wide dynamic range.

For the construction of reagentless enzyme electrodes (e.g., electrodesanalogous to those of Heller et al. and Aizawa et al.) one has to focuson a technique for the modification and functionalization of electrodeand even micro-electrode surfaces to allow the strong binding of theenzyme and the redox mediator taking into account the presuppositionsfor an effective and fast electron transfer between the enzyme and theelectrode. These features requirements are in principle met with enzymeelectrodes based on redox-sensitive hydrogels, however, the manualdeposition of these hydrogels is not compatible with mass-productiontechniques.

The electrochemical deposition of conducting-polymer layers occursexclusively on the electrode surface and can hence be used for theimmobilization of enzymes either covalently using functionalities on thepolymer film or physically entrapped within the growing polymer film. Asthe conducting-polymer film itself does not participate in the electrontransfer, mediator-modified enzymes entrapped within a polypyrrole layerhave been used for the construction of a reagentless oxidase electrode.

Electrochemical deposition methods of the prior art typically use highcurrent density and voltage potential conditions which destroy theorderly Helmholtz double-layer at the electrode surface (U.S. Pat. No.5,215,631 to Westfall). Resulting disorderly depositions at electrodesurfaces produce random polymer structures which lack orientational andpositional order. Aizawa et al. “wired” PQQ-FDH in their sensors withultrathin conductive polypyrrole (PP) membrane as a molecular interface.Electrochemical synthesis of molecular-interfaced FDH on Pt electrodewas prepared by the following two steps: (1) potential-controlledadsorption of FDH, and (2) electrochemical polymerization ofpolypyrrole. These steps employ high voltage and current densityelectrochemical deposition conditions to produce polymer (FDH andpolypyrrole) depositions on the Pt electrode that are randomly oriented.Therefore, this device must operate at high (˜400 mV) operatingpotential resulting in possible interfering cooxidizable species.

What is needed is an improved technique for depositing molecularrecognition groups and associated wiring, if necessary, that provides astrong direct connection between an electrode and the molecularrecognition groups, and allows the molecular recognition groups to bealigned in a common orientation.

SUMMARY OF THE INVENTION

In one aspect, the present invention provides a sensor for sensing thepresence of an analyte component without relying on redox mediators.This sensor may be characterized as including the following elements:(a) a plurality of conductive polymer strands each having at least afirst end and a second end and each aligned in a substantially commonorientation; (b) a plurality of molecular recognition headgroups havingan affinity for the analyte component and being attached to the firstends of the conductive polymer strands; and (c) an electrode substrateattached to the conductive polymer strands at the second ends.

The polymer strands in a common orientation resemble liquid crystals.Preferably, the strands are oriented substantially orthogonal to theelectrode substrate. The conductive polymer strands may be, for example,one or more of multi-stranded nucleic acids, electron transportproteins, synthetic organic and inorganic conducting polymers, metalcrystallite molecular wires, and Langmuir-Blodgett conducting films. Ina particularly preferred embodiment, the conductive polymer strands aredouble-stranded DNA strands.

The headgroup may participate in a redox reaction when contacting amolecule of the analyte component. When this is the case, a mobilecharge carrier is transferred directly to a conductive polymer strandattached to the headgroup, without participating in a redox reaction inthe polymer strand. In one embodiment, the molecular recognitionheadgroups participate in the redox reaction by catalyzing a chemicaltransformation of the analyte component. Examples of such headgroupsinclude oxidoreductases and catalytic antibodies. In one specificexample used repeatedly in this specification, the headgroup is glucoseoxidase.

The sensor headgroups may be chemically homogeneous (e.g., they are allglucose oxidase) or chemically inhomogeneous (e.g., they include amixture of glucose oxidase, cholesterol oxidase, and cholesterolesterase). In one preferred embodiment, when the headgroups areinhomogeneous, the sensor includes a first region on the electrodesubstrate where a first group of chemically homogeneous molecularrecognition headgroups is located and second region on the electrodesubstrate where a second group of chemically homogeneous molecularrecognition headgroups is located. The first and second regions may beseparately addressable so that information signal from the two regionsmay be separately processed and able to indicate whether cholesterol,glucose, or both cholesterol and glucose are present in the analyte forexample.

The electrode substrate should be capable of reporting to an electroniccircuit reception of mobile charge carriers from the conductive polymerstrands. In one specific embodiment, the electrode substrate is a diodesuch as a photovoltaic diode. More generally, the substrate may be adevice element of a device on semiconductor chip (e.g., a gate on anFET).

In a variation of this aspect of the invention, a sensor is provided todetect the presence of a nucleic acid sequence (at a crime scene forexample). The sensor includes (a) a plurality of sequence-specificsingle-stranded nonconductive nucleic acid wires each having at least afirst end and a second end; and (b) an electrode substrate attached tosequence-specific single-stranded nonconductive nucleic acid strands atthe second ends and capable of reporting to an electronic circuit,reception of mobile charge carriers originating from complementarymulti-stranded nucleic acid strands. In this embodiment, when the sensoris exposed to an analyte having the complementary nucleic acid sequence,at least some of the affixed single-stranded nonconductive nucleic acidwires hybridize or anneal with the analyte to form conductivemulti-stranded nucleic acid strands. Thus, charge carriers can betransported to the electrode substrate for detection. In one embodiment,the plurality of sequence-specific single-stranded nonconductive nucleicacid strands are attached to molecular recognition headgroups such thatmobile charge carriers are transferred directly through only annealedmulti-stranded nucleic acid strands when a redox reaction occurs at theattached molecular recognition headgroups.

Another aspect of the invention provides method of detecting aconcentration of an analyte component in an analyte with a sensor havinga structure as described above. The method may be characterized asincluding the following steps: (a) contacting the molecular recognitionheadgroups with the analyte; and (b) determining whether electrons havebeen transferred to the electrode substrate resulting from electronsgenerated by the redox reaction and transferred by the conductivepolymer strands to the electrode substrate. When the redox reactionoccurs at a headgroup, a mobile charge carrier is transferred directlyto a conductive polymer strand attached to the headgroup, without redoxreaction in the polymer strand. The method may further involve (c)monitoring a change in an electronic circuit connected to the electrodesubstrate, the change resulting from reception of mobile charge carriersfrom the conductive polymer strands; and (d) correlating the change inthe electronic circuit with the concentration of the analyte component.

Another important aspect of the claimed invention is a sensor employinga diode, preferably a photodiode. Sensors in accordance with this aspectof the invention may be characterized as including the followingfeatures: (a) a plurality of molecular recognition headgroups having anaffinity for the analyte component and participating in a redox reactionwhen contacting a molecule of the analyte component such that when theredox reaction occurs at a headgroup, a mobile charge carrier isgenerated; (b) a diode having a first electrode to which the pluralityof molecular recognition headgroups are affixed such that mobile chargecarriers generated by the redox reaction are transferred to the firstelectrode; and (c) a circuit for detecting when the mobile chargecarriers are transferred to the first electrode. In a preferredembodiment, the plurality of molecular recognition headgroups areattached to a p-type side of the diode. Also the diode may be a deviceon semiconductor chip including a plurality of devices.

In a further preferred embodiment, the headgroups are attached throughconductive polymer strands arranged as described in the aboveembodiments. Thus, for example, the conductive polymer strands may besubstantially commonly oriented (e.g., orthogonal to the diode surface).

A diode sensor as described above may be used according to a method asfollows: (a) contacting the molecular recognition headgroups with theanalyte; (b) specifying a baseline electrical signal that is presentwhen (i) a stimulus is provided to the diode and (ii) the plurality ofmolecular recognition headgroups are substantially free of the analytecomponent; and (c) detecting a deviation from the baseline electricalsignal, which deviation results from transfer of the mobile chargecarriers to the first electrode when the analyte component comes incontact with the molecular recognition headgroups. The method mayfurther include (d) determining an amplitude of the deviation; and (e)determining an analyte component concentration directly from theamplitude of the deviation without the use of any other information fromthe electrical signal. It has been found that the analyte componentconcentration is sometimes proportional to the amplitude of thisdeviation. Depending upon the type of signal detector employed, thebaseline electrical signal and the deviation from the baselineelectrical signal may be measures of voltage or electrical current.Preferably, though not necessarily, the diode is a photovoltaic diodeand the stimulus provided in the specifying a baseline electrical signalis radiant energy.

Yet another aspect of the present invention is method of forming asensor capable of sensing the presence of an analyte component. Thismethod may be characterized as including the following: (a) contacting asensor substrate (e.g., a device element of a device on semiconductorchip) with a first medium containing mobile conductive polymer strandsor precursors of the conductive polymer strands; (b) applying a firstpotential to the substrate sufficient to form a first structure havingthe conductive polymer strands affixed to the substrate; (c) contactingthe sensor substrate, with affixed conductive polymer strands, with asecond medium containing mobile molecular recognition headgroups; and(d) applying a second potential to the substrate sufficient to affix themolecular recognition headgroups to the affixed conductive polymerstrands. This process produces a sensor structure in which the substrateaffixed to the conductive polymer strands and the molecular recognitionheadgroups also affixed to the conductive polymer strands.

Preferably, the step of applying a first potential is performed at apotential which causes the affixed conductive polymer strands to beoriented in a substantially common direction. This potential may bebetween about 0.001 and 500 mV, for example. The step of applying asecond potential is preferably performed at a potential which causes theaffixed molecular recognition headgroups to be oriented in asubstantially common direction. This second potential may be betweenabout 0.001 and 500 mV. Preferably, though not necessarily, the firstmedium is removed from the sensor substrate following the step ofapplying a first potential. In an alternative embodiment, the secondmedium is obtained from the first medium by performing the step ofapplying a first potential.

If a sensor having separated regions of different headgroups is to becreated, the method may also require isolating a region of the sensorsubstrate prior to the step of contacting the sensor substrate with asecond medium, such that the molecular recognition headgroups aredeposited only in the isolated region. To produce multiple headgroupregions, the steps of isolating a region, contacting the sensorsubstrate with a second medium, and applying a second potential to thesubstrate are performed a second time. The step of contacting the sensorsubstrate with a second medium for a second time employs a secondmolecular recognition headgroup, to form a structure having a firstregion on the sensor substrate having a first group of chemicallyhomogeneous molecular recognition headgroups and a second region on thesensor substrate having a second group of chemically homogeneousmolecular recognition headgroups.

Sensors of this invention provide analyte concentration readings, fastresponses, high sensitivity, high dynamic range, and few erroneousreadings. In a glucose sensor of this invention, glucose concentrationis accurately read despite changes in partial pressure of O₂,atmosphere, altitude, humidity, or sample application of blood.Specifically, the direct wired enzyme sensors of the present inventionovercome the difficulty caused by molecular oxygen reoxidizing a reducedenzyme before that enzyme (or more precisely its redox center) canrelease electrons to the electrode. This is because the directly wiredsensors of this invention may provide electron transfer rates manyorders of magnitude faster than enzymatic reaction rates, and electrontransfer rates of diffusional redox mediators such as O₂ and otherartificial mediators. This provides sub-millisecond digital output fromthe sensing chip.

Chips based on device molecular transistors may be reusable, disposable,reagentless, membraneless. Further, they are amenable to miniaturizationand mass production, do not require complicated three electrode systems(i.e., no working, counter, or reference electrodes) and associatedelectrochemical instrumentation (i.e., no galvinostat or potentiostat),and provide real-time digital output directly from the chip.

Another aspect of the invention pertains to sensors which may becharacterized by the following features: (a) a plurality of conductivepolymer strands each having at least a first end and a second end, andbeing in contact with a molecular insulator that isolates one or more ofthe conductive polymer stands from one or more other conductive polymerstrands on the sensor; (b) a plurality of molecular recognitionheadgroups having an affinity for the analyte component and beingattached to the conductive polymer strands such that when the analyteinteracts with the molecular recognition headgroup one or more mobilecharge carriers are transferred to a conductive polymer strand attachedto the headgroup; and (c) an electrode substrate attached to theconductive polymer strands at the second ends configured to report to anelectronic circuit reception of mobile charge carriers from theconductive polymer strands whereby the presence of the analyte componentis sensed.

In certain embodiments, the plurality of molecular recognitionheadgroups are chemically inhomogeneous. In such embodiments, the sensormay include a first region on the electrode substrate where a firstgroup of chemically homogeneous molecular recognition headgroups islocated and second region on the electrode substrate where a secondgroup of chemically homogeneous molecular recognition headgroups islocated, and wherein the first and second regions are separatelyaddressable. For example, in one region, the molecular recognitionheadgroups may include at least one of glucose oxidase and in anotherregion they may include glucose dehydrogenase.

Another aspect of the invention pertains to a method of forming asensor, which may involve the following operations: (a) affixingconductive polymer strands or precursors of the conductive polymerstrands to a sensor substrate; (b) affixing molecular recognitionheadgroups to the affixed conductive polymer strands, whereby a sensorstructure is formed having the substrate affixed to the conductivepolymer strands and the molecular recognition headgroups also affixed tothe conductive polymer strands; and (c) contacting at least theconductive polymer strands with a molecular insulator that isolates oneor more of the conductive polymer stands from one or more otherconductive polymer strands on the sensor.

In certain embodiments, the method also involves isolating a region ofthe sensor substrate prior to affixing the molecular recognitionheadgroups to the affixed conductive polymer strands, such that themolecular recognition headgroups are deposited only in the isolatedregion.

Still another aspect of the invention pertains to methods of sensing thepresence of an analyte component in an analyte with a sensor including(i) a plurality of conductive polymer strands each having at least afirst end and a second end, and being in contact with a molecularinsulator that isolates one or more of the conductive polymer standsfrom one or more other conductive polymer strands on the sensor, (ii) aplurality of molecular recognition headgroups having an affinity for theanalyte component and being attached to the conductive polymer strandssuch that when the analyte interacts with the molecular recognitionheadgroup one or more mobile charge carriers are transferred to aconductive polymer strand attached to the headgroup, and (iii) anelectrode substrate attached to the conductive polymer strands at thesecond ends. The method involves the following operations: (a)contacting the molecular recognition headgroups with the analyte; and(b) determining whether mobile charge carriers have been transferred tothe electrode substrate resulting from mobile charge carrierstransferred by the conductive polymer strands, at least some of whichare isolated from one another by the molecular insulator, to theelectrode substrate to thereby sense the presence of the analytecomponent.

These and other features and advantages of the present invention will bedescribed in more detail below with reference to the drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a representation of the mechanisms employed in a conventionalredox mediator based biosensor.

FIG. 2 is a representation of a sensor-solution interface in accordancewith this invention and showing a substrate, molecular wire, andmolecular recognition headgroup.

FIG. 3 is a schematic illustration of photodiode sensor in accordancewith an embodiment of the present invention.

FIG. 4A is representation of an electrodeposition step for attachingmolecular wires to a substrate in accordance with an embodiment of thisinvention.

FIG. 4B is representation of an electrodeposition step for attachingmolecular recognition headgroups to molecular wires (deposited as shownin FIG. 4A) in accordance with an embodiment of this invention.

FIG. 5 is a graph showing a current signal generated when glucose iscontacted with a photodiode type GOD glucose sensor in accordance withone embodiment of this invention.

FIG. 6 is a graph showing current and voltage signals generated when thesensor employed in FIG. 5 is subjected to a regimen including contactwith glucose, washing, open circuit, and recontact with glucose.

FIG. 7 is a graph showing current and voltage signals generated from asensor employing GDH on a photodiode when exposed to glucose.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

I. Overview

II. Solid Substrate

III. Sequential Electrochemical and Chemical Deposition Techniques

-   -   A. Electrochemical Atomic Layer Epitaxy (ECALE)    -   B. Sequential Monolayer Electrodeposition (SMED)    -   C. Thin Film Chemical Deposition (CD)    -   D. Electrochemical Molecular Layer Epitaxy (EMOLE)        -   1. Deposition of Uniaxially Oriented Liquid Crystal            Conducting Biopolymers (Proteins and DNA)            IV. Conducting Polymers and Thin Films    -   A. Electron Transport Proteins    -   B. DNA Quantum Wires        V. Molecular Recognition Surfaces    -   A. Oxidoreductases (Redox Enzymes)    -   B. Immunoglobulins        VI. Conduction Mechanisms through Polymers on Solid Substrates    -   A. Energy Bands in Uniaxially Oriented Liquid Crystal Conducting        Biopolymers (Proteins and DNA) and Semiconductor Substrates    -   B. Superconductivity        VII. Applications        VIII. Screening and Assays        IX. Examples

I. OVERVIEW

The present invention relates to sensors, sensor fabrication processesand semiconductor devices that include the sensors. The sensors andrelated devices may be used for recognizing the presence of,quantitating the amount of, and/or continuously monitoring the level of,one or more selected components in a solid, semi-solid, liquid, or gasmixture. Preferably, an active molecular recognition surface is “hardwired” to the substrate surface (e.g., a semiconductor surface) by anoriented liquid crystal wire that is itself conductive. The molecularrecognition surface may be of biologically active material of the typeconventionally employed in sensors. The substrate may be patterned orunpatterned and may include (particularly when semiconductors areinvolved) a conductive coating such as a metal between the underlyingbulk substrate and the liquid crystal wire.

Hard wiring as that term is used herein may be achieved, in oneembodiment, via electrochemical fabrication methods described in detailbelow. Generally, such methods make use of low-cost, rapid-prototypingsequential electrochemical and chemical deposition techniques such aselectrochemical molecular layer epitaxy (EMOLE) which perform “molecularwiring” and “molecular soldering” procedures. The liquid crystal wiringarrangement preferably provides a “lawn” of commonly oriented “moleculardevices” each including a single molecular recognition site “headgroup”and attached molecular wire “tail.” For context, each such device mightrange in size from about ˜2 to 2500 Å² surface area (e.g., enzyme,enzyme co-factor, substrate, supramolecular assembly, cavitand,host-guest complex, ligand, receptor, antibody, antigen, etc.).

Biosensors of the present invention may require very low operatingpotentials. In a preferred embodiment, extended conformation of straightuniaxially oriented liquid crystal DNA wires are stuck into the GODactive site/redox center of the prosthetic group FAD/FADH₂, to providean electron transfer pathway to the surface of a p-n homojunctionsemiconductor solar cell substrate. A pair of electrons per enzymeturnover event injected from the wires combine with a pair of holes inthe p-type semiconductor layer, interfering with the normal photocurrent(i.e., electron/hole pair recombination) occurring in the solar cell.The oriented liquid crystal enzyme (molecular recognition headgroup) andattached oriented liquid crystal DNA wire tail constitute a moleculartransistor. The device communicates with a solid substrate (i.e., p-nhomojunction) through the uniaxially oriented liquid crystal DNA wiretail interconnects. One end of the DNA wire is stuck in the orientedliquid crystal enzyme active site/redox center and the other end isstuck into the p-type semiconductor layer providing a direct connectionbetween the protein enzyme, DNA, and semiconductor substrate.

In general, the sensors of this invention may be categorized based upontheir transduction and/or gating mechanisms of the headgroup(s):switched or gated by optical (optoelectronic), chemical(chemoelectronic), magnetic (magnetoelectronic), radioactive(radioelectronic), thermal (thermoelectronic), mechanical(piezoelectronic), or electrical (voltage, current, resistivity,capacitance).

FIGS. 2 and 3 depict sensors structures in accordance with certainpreferred embodiments of the present invention. FIG. 2 presents across-sectional view of a surface region of a sensor 12. As shown,sensor 12 includes an electrode 14 which is preferably made from siliconor another semiconductor substrate. Attached to electrode 14 is aplurality of conducting polymer strands 16. In a preferred embodiment,each strand is a DNA double-stranded molecule. Conductive polymerstrands 16 are orientated substantially in a common direction which isshown to be normal (orthogonal) to substrate 14. Strands 16 are coupledto substrate 14 in a manner that allows direct electrical influencebetween these two features in the sensor. For example, the connectionmight allow electrons to be directly transferred from strands 16 tosubstrate 14 so that circuitry coupled to substrate 14 can detectinjection of electrons. In addition, a potential applied to substrate 14may influence the physical state of conductive polymer strands 16.

As will be described in more detail below, a preferred process foraffixing polymer strands 16 to substrate 14 provides this directelectronic coupling and in addition orients the strands 16 along asubstantially common axis. Because strands 16 are oriented in asubstantially common direction, they will sometimes be collectivelycharacterized herein as a liquid crystal.

Note that liquid crystal conductive polymer strands such as those shownin FIG. 2 take the form of a “lawn” having first ends attached tomolecular recognition headgroups 18 and second ends attached toelectrode 14. As will be described below, headgroups 18 may take manydifferent forms. Generally, they should change physical or chemicalstate in response to the presence of a particular component in analyte20. In a preferred embodiment, molecular recognition headgroups 18 areenzymes which undergo a redox transformation in response to contact witha specified analyte component. For example, the analyte may include aligand or substrate component 25 which selectively binds with and ischemically modified by headgroups 18. Preferably, the chemicalmodification is accompanied by generation of electrons which candirectly transferred to strands 16 and from there to electrode 14.Depending upon the type of molecular recognition headgroup 18 employedin the sensor 12, the thickness of a headgroup layer on top of theconductive polymer lawn 16 may be between about 5 and 150 angstroms.

Importantly, no mediator is required in this sensor design, so electrontransfer is direct and fast from headgroup 18 to electrode 14. Further,because the polymer strands 16 are commonly oriented, headgroups 18 areoptimally presented for sensing the desired analyte component. That is,headgroups 18 are not sterically hindered by polymer strands 16 or otherstructures.

While the plurality of conductive polymer strands 16 may have a ratheruniform length as depicted in FIG. 2, this need not be the case. Morefrequently, the individual polymer strands will have a wide range oflengths. This will be due to inherent variations in polymerizationtechniques or the polymer shearing techniques. Of course, thedistribution of polymer strand lengths can be made more uniform bypassing a raw collection of polymer strands through a chromatographycolumn, electrophoretic gel, ultrafiltration membrane, or other sizingapparatus. In a preferred embodiment, the average strand length ofconductive polymer strand 16 is between about 2 and 1,000 Å. Morepreferably, the length is between about 10 and 100 Å, and mostpreferably between about 3 and 40 Å. When DNA is employed as theconductive strands, the width of the individual sensor strands is in theneighborhood of 20 Å.

In a preferred embodiment, the substrate 14 is a p-type electrode of asilicon photodiode. It may include, though this is not always necessary,a metallic back plate 22 for providing an ohmic contact between polymerstrands 16 and bulk silicon electrode 14. Such back metal plates areconventionally used in semiconductor devices as terminals for connectionto an external circuit. The back metal plate 22 may be made from anysuitable conductive metal or alloy, including but not limited toaluminum, copper, silver, gold, and platinum. Region 24 represents theclose packed liquid crystal spacing between EMOLE deposited molecularrecognition headgroups. Molecular recognition headgroups whosedimensions are greater than the width of underlying molecular wires towhich they are attached occupy region 24.

In a preferred embodiment, the semiconductor substrate forms part of arectifying diode such as a photodiode. FIG. 3 provides a schematicillustration of a photodiode based biosensor in accordance with oneembodiment of the present invention. A sensor 50 includes a photodiode52 including an n-type region 53 and a p-type region 54. Generally, anyconventional photodiode may be employed with this invention, but itshould have a surface suitable for affixing conductive polymer strandsand molecular recognition headgroups as described above. To this end,p-type region 54 may be provided with or without a back metal ohmiccontact 56 as shown. A plurality of strands of conductive polymer 58 areaffixed at one end to back-metal plate 56. The other ends of polymerstrands 58 are attached to a collection of molecular headgroups 62. Theresulting structure, as illustrated, may be identical with the structureof elements 14, 22, 16 and 18 as shown in FIG. 2.

Photodiode 52 includes a depletion region 60 which automatically formsat the p-n semiconductor junction. As is known to those of skill in theart, depletion regions form at these interfaces because mobile holesdiffuse from p-type regions into n-type regions just across theinterface where they are combined with electrons available in the n-typeregion. Similarly, mobile electrons in the n-type region diffuse acrossthe interface to the p-type region where they combine with holes. As aresult, within the reach of charge carrier diffusion, essentially allmobile charge carriers are depleted.

When light (or other radiant energy of appropriate wavelength) is shownon a photodiode such as photodiode 52, some holes and electrons crossthe semiconductor band gap and provide additional mobile charge carrierswhich can be drawn out of photodiode 52 by an applied potential orexternal short circuit connection. Applied potentials or external shortcircuit connections may be made through a digital multi-meter 64, avariable potential power supply, a battery, another photodiode, or apotentiostat, for example. Of course, many other potential sources orexternal short circuit connections may be employed. A multi-meter 64 hasthe advantage of being inexpensive yet able to detect the amount ofcurrent flowing as a result of the incident light. Additional electronsare attracted to p-type region 54 by the excess holes generated by thelight. Similarly, electrons flow out of n-type region 53 because thereare now excess electrons by virtue of the light excitation. This currentflows through a line 66, multi-meter 64, and a line 68. Note that line68 is electrically connected to back plate 56. Similarly, line 66 isconnected to a metal back plate 70.

When electrons are injected into the p-type region 54, they may combinewith and thereby annihilate holes. Thus, the photocurrent amplitude isreduced. Detection of this deviation from normal photocurrent specifiesthat an analyte component has been detected. It has been found that theamplitude of this deviation is proportional to the analyte componentconcentration. Further, it has been found that the deviation is presentin both the current and voltage associated with the photodiode.

It should be understood that the sensors of this embodiment of theinvention can be formed on any type of diode in which an externalstimulus generates a baseline current. Such stimulus may be heat(thermally generated charge carriers), electric field, radiation, etc.In each case the baseline current is at least partially “quenched” byelectrons or holes injected from the lawn of molecular devices when aspecified analyte component is present. Amplitude of the deviation frombaseline is often proportional to concentration of the analytecomponent. A simple calibration curve for each chip can be used todetermine concentration of the analyte component(s) in unknown samples.

In a particularly preferred embodiment, the sensor is divided into aplurality of regions, each capable of sensing the presence of adifferent analyte component. For example, a first region might include,as molecular recognition headgroup, glucose oxidase to sense thepresence of glucose, a second region might include cholesterol esteraseand cholesterol oxidase to sense the presence of cholesterol, a thirdregion might include alcohol dehydrogenase to sense the presence ofethanol, etc. Each of these regions will be separately addressable byelectronic circuitry to uniquely identify the presence a particularanalyte component. Each of the sensor regions could be made separatelyaddressable by specialized circuitry employed in conventional integratedcircuits. While the circuitry need not be particularly complex, suchdevices allow very sophisticated processing of the data provided by thesensor regions.

In certain embodiments, a molecular plastic insulator is deposited onand/or around a defined sensor region. In some situations, the insulatoris deposited by a process that follows EMOLE fabrication of the sensorregion. As indicated, this may be a macro-, micro-, and/or nanometerscale sensor region. The molecular plastic insulator provides anelectrical insulation sheath around a bundle molecular wires and/orindividual molecular wires that make up or define a specific macro-,micro-, and/or nano-sensor region. In certain embodiments, the definedsensor regions are separately addressable. The molecular insulatorminimizes and/or prevents electrical (i.e., electron and hole)cross-talk between separately addressable sensor regions.

After EMOLE fabrication of a specific sensor region, a liquid molecularplastic insulator may be deposited onto the surface of the sensorregion, and allowed to remain in contact with the sensor region andundergo a physical and/or chemical reaction for a period of time (e.g.,1 millisecond to 1 week, typically one second to 20 hours). Thereafter,any excess molecular plastic insulator may be washed off withappropriate miscible solvent(s).

Examples of molecular plastic insulators suitable for the describedpreferred embodiment are general classes of hydrophobic amines such asquaternary amines, tertiary amines, secondary amines, primary amines,and a myriad of other alkyl cationic and anionic molecular structures.Some specific examples fall into broad classes of polymers and monomers,each of which bind to conducting molecular wires on the sensor to resultin insulating sheaths on the molecular wires.

Molecular plastic insulators bind by, e.g., covalent, non-covalent,physisorbed, chemisorbed, ionic, hydrophobic, hydrophilic, and van derWaals interaction binding to particular conducting biopolymer molecularwires and provide a fixed, large band gap electrical insulator toconducting biopolymers (e.g., proteins and DNA). The molecular plasticinsulator provides an electrical insulation sheath around a bundle ofconducting biopolymer molecular wires and/or individual molecular wiresthat make up or define a specific macro-, micro-, and/or nano-sensorregion.

DNA, RNA, PNA, and other nucleic acid and protein binding molecules can(i) intercalate into multi- and single-strand nucleic acids, (ii) canbind to single-stranded DNA, (iii) can bind to single-stranded RNA, (iv)can bind to the major groove of multi- and single-strand nucleic acids,(v) can bind to the minor groove of multi- and single-strand nucleicacids, (vi) can bind to poly-anion backbone of multi- andsingle-stranded nucleic acids and proteins, (vii) can bind topoly-cation backbone of multi- and single-stranded nucleic acids andproteins; (viii) can form salt-bridges with multi- and single-strandednucleic acids and proteins; (ix) can form hydrogen bond (H-bond) basepairs to multi- and single-stranded nucleic acids and proteins, (x) canbind to sequence-specific sites of multi- and single-stranded nucleicacids and proteins, (xi) can bind to non-sequence specific sites ofmulti- and single-stranded nucleic acids and proteins, (xii) can formhydrophobic and hydrophilic interaction bonds to nucleic acids andproteins, etc. Examples of DNA, RNA, PNA, and other nucleic acid andprotein binding molecules include but are not limited to the followingexamples.

a. DNA binding proteins

b. RNA binding proteins

c. PNA binding proteins

d. DNA binding peptides

e. RNA binding peptides

f. PNA binding peptides

g. DNA polymerases

h. RNA polymerases

i. Restriction endonucleases

j. Operator binding proteins

k. Gene regulatory binding proteins

l. Promoter binding proteins

m. Spermine and analogs

n. Sperimidine and analogs

o. Amine-derivatized dendrimers and analogs

p. Histones and analogs

q. Polyamines and analogs

r. Aromatic ring organic molecule nucleic acid binders

s. Aromatic ring organo-metallic molecule nucleic acid binders

t. Aromatic ring metallic molecule nucleic acid binders

u. Poly-cation organic molecule nucleic acid binders

v. Poly-cation organo-metallic molecule nucleic acid binders

w. Poly-cation metallic molecule nucleic acid binders

x. Sequence-specific nucleic acid binding molecules

y. Non-sequence specific nucleic acid binding molecules

z. Anti-sense molecules

aa. RNAi molecules

bb. MicroRNAs

cc. Antibody molecules and fragments

dd. Small molecule protein binders

ee. Large molecule protein binders

ff. Protein and peptide molecule protein binders

gg. Hydrophobic interaction molecule protein binders

hh. Hydrophilic interaction molecule protein binders

ii. Ionic interaction molecule protein binders

jj. Protein bioconjugates

kk. Nucleic acid bioconjugates

ll. van der Waals protein binders

mm. van der Waals nucleic acid binders

The molecular devices (headgroup and conductive strand affixed to anelectrode surface) in each region may be formed by processes similar tothose employed in integrated circuit fabrication. For example, certainregions could be exposed to light radiation shown through a patternedreticle. Those regions would be selectively activated or protecteddepending upon the use of appropriate chemical protecting groups. Aliquid crystal conductive polymer region or headgroup region would thenbe formed on the reactive regions. Such processes are described in U.S.Pat. No. 5,252,743 issued to Barrett et al. and Pritchard et al.,“Micron-Scale Patterning of Biological Molecules” Angew. Chem. Int. Ed.Engl., Vol. 34, No. 1, pages 91-93 (1995), for example, which isincorporated herein by reference for all purposes. Alternatively, anelectric potential could be selectively applied to certain of thesubstrate regions to selectively electrodeposit the distinct sensorregions.

II. SOLID SUBSTRATE

Various solid substrates may be employed in the invention. The solidsubstrate should undergo a detectable change in response to anelectrical stimulus from the molecular wire. The substrate material maybe biological, nonbiological, organic, inorganic, or of a combination ofany of these, existing as particles, strands, precipitates, gels,sheets, tubing, spheres, containers, capillaries, pads, slices, films,plates, slides, etc. The substrate may have any convenient shape such asdisc, square, sphere, circle, etc. The substrate and its surfacepreferably, though not necessarily, form a rigid support on which tocarry out the reactions and fabrication processes described herein. Thesubstrate and its surface may also be chosen to provide appropriatecrystal or non-crystal lattice structure, wafer or thin filmorientation, n- and p-type doped materials, surface texture, back metalpattern, grid metal pattern, surface chemistry, etc. The raw macro-solidsubstrate may be composed of a semiconductor or standard electricalcomponent. Preparation of surfaces by lapping, polishing, chemicaltreatment, ion implantation, photolithography, etching, chemical vapordeposition (CVD), molecular beam epitaxy (MBE), etc. may provide apatterned macro-solid substrate suitable for further processing by meansof the present invention.

Various semiconductor substrates may be employed in the invention. Thesemiconductor substrate may be biological (e.g., lipid bilayers,membrances, detergent solubilized membrane fragments containing embeddedprotein electron transport pathways, blood brain barrier (BBB),epithelial linings, intestinal linings, intracellular membranefragments, intracellular organelles, different tissue cell surfacetypes, membrance surfaces from different blood types of red blood cells,membrane surfaces from different types of lymphocytes, macrophages, andwhite blood cells, lyposomes, arterial and venous blood vessel walls,neuronal conduction pathways, etc.), nonbiological, organic, inorganic,or of a combination of any of these. Usually, the semiconductorsubstrate will be composed of silicon, doped diamond, indium tin oxide,tin oxide, gallium arsenide, cadmium sulfide, cadmium selenide, cadmiumtelluride, germanium, copper indium diselenide, copper indium disulfide,copper indium ditelluride, zinc sulfide, zinc selenide, mercurytelluride, mercury selenide, graphite, etc. or combinations thereof.Other substrate materials will be readily apparent to those of skill inthe art upon review of this disclosure. In a preferred embodiment thesemiconductor substrate is a p-n doped polycrystalline ormonocrystalline silicon (e.g., having a surface crystallographicorientation in the <100> or <111> direction) or copper indium diselenidemonocrystalline thin film deposited onto glass.

A semiconductor substrate may form part of a homojunction device wherethe same semiconductor material is employed on either side of the p-njunction, differing only in dopant type; or heterojunction device, wherethe materials on either side of the p-n junction are semiconductors butdifferent semiconductors. Processes and chemistries for homo- andheterojunction device manufacture are known in the art and will not bedescribed in significant detail. A conventional photovoltaic solar cellis an example of a semiconductor homojunction device. It is a standardn-p junction, rectifying diode with contact metallization partiallycovering its emitter to allow light entrance.

In a rectifying diode, for example, conducting back metal contactpatterns may be located on the p-type surface and conducting grid metalcontact patterns may be located on the n-type surface. Such back metalpatterns are generally used for the purpose of providing an ohmiccontact to the semiconductor diode. In the present invention, they maybe used for attaching highly conductive terminal contacts of theconducting polymer to the semiconductor substrate surface in specificregions as described in the next section. Back or grid metal contactsare typically made from a conductive metal layer such as aluminum,copper, gold, silver, etc. The back or grid metal may be textured andmay adopt lattice matching of underlying monocrystalline <100> or <111>silicon surfaces upon which it is deposited. Alternatively, theconducting polymers or thin films of this invention may be directlyconnected to p-type polycrystalline or monocrystalline surfaces, withoutthe need for back metal.

The raw macro-solid substrate may be connected to or comprise standardelectrical components (e.g., transistor, diode, electrode, semiconductorheterojunction, semiconductor homojunction, Schottky barrier, capacitor,resistor, inductor, CMOS, TTL CMOS, FET, ISFET, MOSFET, ENFET, REFET) orcombinations thereof (See e.g., U.S. Pat. No. 5,126,921 to Fujishima etal.; U.S. Pat. No. 5,108,819 to Heller et al.; U.S. Pat. No. 5,403,700to Heller et al.). Memory and logic circuitry on such chips can beemployed to interpret sensor signals. In a preferred embodiment, thesensor wiring will be attached to transistor gates, sources, or drains(to control potential) or to other circuit or device components tocontrol current. Preparation of active surfaces on the semiconductorsubstrate may be accomplished by various fabrication techniquesincluding, for example, lapping, polishing, chemical treatment, ionimplantation, photolithography, etching, chemical vapor deposition(CVD), molecular beam epitaxy (MBE), etc.

It may be possible to wire only few or even one conductive polymerstrand to a device element such as gate of a FET. Using availabletechnology reported by Yoo et al. in Science, entitled “ScanningSingle-Electron Transistor Microscopy: Imaging Individual Charges”, Vol.276, pages 579-582 (1997) (which is incorporated herein by reference forall purposes), source, drain, and gate elements of very small dimensionshave been fabricated on a scanning tunnelling microscope (“STM”) tip.Such devices have been reported to detect transfer of single chargecarriers. By attaching one or a few conductive polymers (and associatedheadgroups) to the gate of such device, for example, a single bindingevent (at single headgroup) could be detected. If the individual devicesare made separately addressable, each polymer strand/headgroupcombination could form a molecular transistor of very small dimensions.Separately addressable STM tips are discussed by Service in Science,“Atomic Landscapes Beckon Chip Makers and Chemists” Vol. 274, pages723-724 (1996).

III. SEQUENTIAL ELECTROCHEMICAL AND CHEMICAL DEPOSITION TECHNIQUES

Sequential electrochemical or chemical deposition techniques may be usedto attach molecular recognition surfaces to conductive polymers and toattach conductive polymers onto semiconductor wafer substrates preparedas described above. Specifically, the present process methods of thisinvention may employ various processes related to electrochemical atomiclayer epitaxy (ECALE), sequential monolayer electrodeposition (SMED),and thin film chemical deposition (CD) in a process referred to hereinas electrochemical molecular layer epitaxy (EMOLE) to deposit,polymerize, and/or orient monomers, polymers, macromolecules, or thinfilms into liquid crystal conducting polymers or “molecular wires” withhighly conductive terminal contacts. Preferably, one terminal contact ofthe formed one-dimensional molecular wire is “molecularly soldered” orelectrically connected to the substrate surface (i.e., the back metalcoated on a p-type surface of the semiconductor homojunction substrate).The other terminal contact is directed outward by virtue of extendedliquid crystal conducting polymer orientation perpendicular to thesubstrate surface as illustrated above in FIG. 2. Repeat of analogousdeposition techniques are used to “molecularly solder” or electricallyconnect an active molecular recognition headgroup to the free terminalcontacts (also illustrated in FIG. 2) permitting rapid and direct chargeconduction from the molecular recognition sites to the semiconductorsubstrate.

In a preferred embodiment of the invention, sequential deposition occursonly in specific regions of the semiconductor substrate (e.g., onspecific electrically or chemically activated surface regions of thesubstrate electrode). This provides a patterened surface of individuallywired molecular recognition sites.

Examples of three sequential deposition techniques (electrochemical andchemical) and their application to production of atomic layers ofcompound semiconductors and conducting polymers are described below inSection III, A-C. A modified form of these processes calledelectrochemical molecular layer epitaxy (EMOLE) may be employed tofabricate a single sensor site or an array of sensor sites.

A. Electrochemical Atomic Layer Epitaxy (ECALE)

The epitaxial growth of semiconductors is an important and active areaof research. The development of new, low temperature techniques for thepreparation of high-quality semiconducting thin-film materials is offundamental importance to the semiconductor chip industry. Considerableeffort has been devoted to study the epitaxial growth of these materialsin vacuum (e.g., molecular beam epitaxy (MBE). Electrodepositionrepresents an alternative to the expense of vacuum techniques. Inaddition, electrochemistry is usually performed near room temperature,and therefore avoids the interdiffusion problems associated with thehigh temperatures used in vacuum deposition methods. Research has beendirected towards the epitaxial electrodeposition of II-VI compoundsemiconductors. A method for epitaxial electrodeposition and digitaletching, electrochemical atomic layer epitaxy (ECALE), is beingdeveloped. The method involves the alternated electrodeposition ofatomic layers of the constituent elements which make up a compound.Deposition is limited to an atomic layer by the use of underpotentialdeposition (UPD). UPD refers to a surface-limited process whereby adepositing element forms a compound with substrate surface atoms at apotential below that required for bulk deposition of the element.Deposition of the element proceeds until the surface is “covered”. Afterthe surface is covered, subsequent deposition requires a higherpotential to promote bulk deposition. Thus, UPD is usually limited tomonolayer coverage.

ALE (atomic layer epitaxy) refers to a series of vacuum based methodsfor semiconductor growth where a compound is formed a monolayer at atime by the alternated deposition of atomic layers of the constituentelements. ALE is applicable to a variety of thin film formation methodssuch as molecular-beam epitaxy (MBE), metalloorganic molecular beamepitaxy (MOMBE), chemical vapor deposition (CVD), metalloorganicchemical vapor deposition (MOCVD), etc. These vacuum methods involvesuch problems as the need for careful control of reactant fluxes inorder to obtain epitaxial deposits. ALE is currently under developmentwhich allows less stringent control of growth parameters. Unique to ALEis compound growth of one atomic layer at a time. This technique relieson surface-specific reactions which result in only a monolayer ofreactivity. If the reactant is an elemental vapor, the substancetemperature is adjusted so that bulk deposits sublime while the firstmonolayer remains due to an enhanced stability resulting from compoundformation. After pumping (evacuation) of the first element, a similarprocedure is performed with the second element. For a compound such asCdTe, a layer of Cd is formed followed by a layer of Te. Thin filmgrowth is achieved by repeating the cycle.

In the formation of a compound such as GaAs by ALE in the MOCVD mode, aflux of H₃As, an arsenic precursor gas, is exposed to the substrate at atemperature which allows formation of a single As surface layer. Allexcess H₃As is subsequently pumped away under high vacuum. The As atomiclayer is stabilized by compound formation with previously deposited Ga.A flux of tetramethyl gallium (TMG), a gallium precursor gas, is thenexposed to the surface, and similarly an atomic layer of Ga is formed.Excess gas is pumped away under high vacuum. Thin films are produced byrepeating this cycle.

ECALE is the electrochemical analog of atomic layer epitaxy (ALE)employing UDP in place of temperature control to deposit monolayers. Useof UPD in order to electrodeposit atomic layers of both elements, atpresent, requires that one element be deposited by reductive UPD whilethe other is deposited by oxidative UPD. In this way, oneunderpotentially deposited element can be held on the surface at thepotential used subsequently to deposit the other element. In theformation of a compound such as CdTe, Te can be oxidativelyunderpotentially deposited from Te²⁻ at a fairly negative potential.Cadmium can next be reductively underpotentially deposited from a Cd²⁺solution at a more positive potential, where previously deposited Teremains stable. Electrodeposited semiconductors do not have to beannealed as in ALE which is typically done for 15 minutes at 300° C.

Digital etching, the reverse process of deposition, is a naturalextension of the ECALE method. Increasing the negative voltage potentialto strip or etch monolayers is possible. A method for the digitalelectrochemical etching of compound semiconductors in an electrochemicalflow cell system in which alternating electrochemical potentials areapplied between a reference electrode and the compound semiconductorsufficient to strip portions, preferably atomic layers, of the elementsof compound semiconductors from the compound semiconductors is describedin Stickney et al.: U.S. Pat. No. 5,385,651 and Stickney et al.: WO94/28203.

B. Sequential Monolayer Electrodeposition (SMED)

Sequential Monolayer Electrodeposition (SMED) provides monolayers ofII-VI compound semiconductors and is related to the ECALE methoddescribed above. However, unlike the ECALE method, all depositedelements are provided in the same electroplating solution. They arecodeposited and then one which deposited in excess is electrochemicallystripped away. For example, Cd²⁺ and Se²⁻ may be deposited from the sameelectroplating solution by cyclic voltammetric deposition at fast scanrates with a nickel rotating disk electrode. The procedure was designedto eliminate the problem of bulk Se formation, using a cyclic depositionscheme that cathodically deposits submonolayer amounts of CdSe and alarge stoichiometric excess of Cd. The excess Cd is then stripped off bysweeping the electrode to a positive potential as part of thevoltammetry cycle (Cd is readily stripped close to its thermodynamicreduction potential). Since the CdSe phase has a large negative freeenergy of formation (ΔG°_(f,298K)=−141.5 kJ mol⁻¹), it was thought thatany free Se that is deposited in this process will react with the excessCd to form CdSe and not lead to large amounts of excess Se in the film.The net result is thus the sequential deposition of stoichiometric CdSea monolayer (or less) at a time. It has been reported that such aprocedure leads to compositionally homogeneous, stoichiometric films andmay be a general method to electrodeposit binary materials with largethermodynamic or kinetic stabilities. (Kressin, A M; Doan, V V; Klein, JD; Sailor, M J: “Synthesis of Stoichiometric Cadmium Selenide Films ViaSequential Monolayer Electrodeposition” Chem. Mater. 3(6): 1015-1020,1991).

C. Thin Film Chemical Deposition (CD)

Conducting polymers continue to look promising as the active elements ofelectronic and chemical devices such as flexible light-emitting diodes,chemical sensors and photovoltaic devices. As a result, the thin filmprocessing techniques for these materials have become increasinglyimportant to the successful fabrication and optimization of usefulall-organic thin film devices. Techniques such as spin coating,electrochemical deposition, and Langmuir-Blodgett thin film transferhave all been utilized with varying degrees of success to manipulateconjugated polymers into thin films. Fou et al. (Fou, A C; Ellis, D L;Rubner, M F: Molecular-Level Control in the Deposition of UltrathinFilms of Highly Conductive, In-Situ Polymerized P-Doped ConjugatedPolymers. Mater. Res. Soc. Symp. Proc. 328:113-118, 1994.) has describeda thin film processing technique that has been developed for thefabrication of ultrathin films of conducting polymers withangstrom-level control over thickness and multilayer architecture.Molecular self-assembly of in-situ polymerized conjugated polymersconsists of a layer-by-layer process in which a substrate is alternatelydipped into a solution of a p-doped conducting polymer (e.g.,polypyrrole, polyaniline) and a solution of a polyanion. In-situoxidative polymerization produces the more highly conductive,underivatized form of the conjugated polymer, which is deposited in asingle layer of precisely controlled thickness (30 to 60 Å). Thethickness of each layer can be fine-tuned by adjusting the dipping timeand the solution chemistry. The surface chemistry of the substrate(i.e., hydrophobic, charged, etc.) also strongly influences thedeposition, thereby making it possible to selectively deposit conductingpolypyrrole onto well defined regions of the substrates.

D. Electrochemical Molecular Layer Epitaxy (EMOLE)

Electrochemical molecular layer epitaxy (EMOLE) is a processingtechnology used to engineer the structure and properties ofmacromolecules deposited on a substrate surface in order to producehighly organized molecular materials. Preferably, this processing yieldsliquid crystal structures of the type described above. Typically,crystallization is viewed as producing homogenous and well orderedmaterials made of one or a few kinds of atoms or small molecules. It isalso possible though to crystallize larger and more complex moleculessuch as proteins, DNA, supramolecular assemblies such as ribosomes, andeven virus particles with atomic masses in excess of 100 milliondaltons. In fact, this is a necessary step in elucidating the structureof many macromolecules. Co-crystallization of two or more differentcomponents is also possible. The present invention provides EMOLEtechniques to produce layers of two-dimensional crystals or generallywell ordered arrangements of interconnected macromolecules for theproduction of a biosensor. EMOLE as described herein generally employslow current density and potential (which maintains the Helmholtzdouble-layer) to deposit uniaxially oriented liquid crystal conductingbiopolymers (proteins and DNA) at substrate surfaces.

Preferably, methods of this invention employ EMOLE to deposit, attach,polymerize, and/or orient monomers, polymers, macromolecules, or thinfilms into liquid crystal conducting polymers or “molecular wires” withconductive terminal contacts. “Thin film” is a term used herein to meana well defined atomic or molecular deposition layer on a flattwo-dimensional substrate. Thin films can be made by many techniques(i.e., ALE, CVD, Langmuir-Blodgett, dip coating, spin coating, EMOLE,etc.) and be composed of many materials. Thin films can sometimes becharacterized as a “lawn” or “liquid crystal.”

Conditions which promote oriented liquid crystal polymers will bepresented below. EMOLE may be employed to form conductive electronicconnections at each end of the oriented liquid crystal conductingpolymers (i.e., the headgroup end and the substrate end). By connectingthem at a first end of conductive polymer strands in a liquid crystalorientation, the molecular recognition headgroups are stericallyunhindered in their chemical or biochemical binding/recognition ofanalyte species. As a consequence of an analyte binding event to amolecular recognition site, rapid electron or hole transfer from theoriented liquid crystal molecular recognition site through the attachedoriented liquid crystal conducting polymer or thin film, to thesemiconductor substrate will produce a signal. The amplitude of thesignal or number of electrons or holes tunneling to the semiconductorsurface taken in aggregate will reflect the amount of specified analytespecies present.

In a preferred embodiment of this invention, a first electrodepositioncycle affixes strands of a conducting polymer on a substrate (e.g., ap-type surface of a semiconductor such as a p-n junction solar celldescribed above). This is depicted in FIG. 4A. In this cycle, a firstmedium 402 containing a polymer 404 to be deposited (or a precursor ofthat polymer such as monomers) is contacted with a substrate 406.Preferably, though not necessarily, medium 402 is a liquid solubilizingthe polymer strands. Medium 402 may be held within a container 407 asshown, or may passed over substrate 406 in a continuous flow reactor. Apotential is then applied to substrate 406 via a circuit 408 to drivethe first cycle and deposit a lawn of immobilized polymer strands 410.Note that circuit 408 includes substrate 406, medium 402, a counterelectrode 412, and a power supply 414. If polymer strands 404 have apositive charge, then a negative potential is applied to the substrate;but if they have a negative charge, a positive charge is applied to thesubstrate. In either event, the potential and/or current density shouldbe controlled to ensure that (1) the polymer is affixed to the substratewith strength to allow electron transport, and (2) the deposited polymerstrands have a substantially common orientation. It may be desirable toinclude a charge group on only one end of polymers 404 so that that endis selectively coupled to the surface of substrate 406. If the polymerstrand is a nucleic acid, the charge group could be attached byincluding it at one end of a nucleic acid strand (designed much like aconventional nucleic acid probe) which strand is complementary to an endof the nucleic acid to be affixed. Of course, other techniques forattaching charge groups (or other functional groups) to one end of apolymer strand are known in the art and may be profitably employed inthe context of the present invention.

In a specific embodiment, electrodeposition current density ranging fromabout 10 to 300 μA cm⁻² and voltage potential ranging from about 10 to300 mV can be generated by light induced photoconduction at the n-typeand p-type surfaces of the submerged solar cell. Deposition cyclevariables include i) applied potentials (i.e., magnetic/voltage); ii)solution condition (i.e., concentration of deposited material, pH,electrolyte, solvent, temperature, etc.); and iii) semiconductorsubstrate (i.e., polycrystalline, monocrystalline, single-crystal faceorientation, smooth or textured surface, metal contact coating, latticematching of coating, etc.). As will be understood to those of skill inthe art, these variables may be adjusted to produce an optimalmolecular-scale structure.

For example, the following guidelines may be employed to depositsuitable molecular wires. First, applied potentials must be low enough(e.g., 0.001 to 1500 mV) to maintain a Helmholtz double-layer duringelectrodeposition of conducting polymers and molecular recognitionheadgroups onto semiconductor substrate. Applied potential ranges willvary depending on the size, charge density, counter ion, and viscosityof the to be deposited material. Second, current densities must be lowenough (e.g., 0.001 to 1500 μA cm⁻²) to maintain a Helmholtzdouble-layer during electrodeposition of conducting polymers andmolecular recognition headgroups onto semiconductor substrate. Currentdensity ranges will vary depending on the size, charge density, counterion, and viscosity of the to be deposited material. Third, thesemiconductor substrate should be chosen to maintain a Helmholtzdouble-layer during electrodeposition of a uniaxially oriented liquidcrystal structure on the surface of the semiconductor substrate. Asnoted, it may be polycrystalline or monocrystalline, having smooth ortextured surface. It may also have a metal contact coating.

Further, the solution conditions should meet certain specific criteria.For example, the concentration of deposited material should be lowenough (e.g., 0.001 to 10 mg/mL) to maintain a Helmholtz double-layerduring electrodeposition of conducting polymers and molecularrecognition headgroups onto semiconductor substrate. Further, the pHshould be adjusted to approximately two (2) pH units above or below thepK_(a) or pI of the conducting polymer or molecular recognitionheadgroup to produce a polymer of opposite charge from the surface ofthe semiconductor substrate. Still further, the electrolyte should bechosen to have a counter ion type and electrolyte concentration (e.g., 0to 150 mM salt) that maintains a Helmholtz double-layer duringelectrodeposition of a uniaxially oriented liquid crystal structure onthe surface of the semiconductor substrate. High electrolyteconcentration will produce too much current and destroy the Helmholtzdouble-layer during electrochemical deposition processing. In addition,the solvent should be chosen from a range of organic and aqueoussolvents and co-solvents to maintain a Helmholtz double-layer duringelectrodeposition of a uniaxially oriented liquid crystal structure onthe surface of the semiconductor substrate. Conducting polymers andmolecular recognition headgroups should be soluble in the solvent orco-solvent used. Finally, the temperature should be greater than thefreezing point (fp) and less than the boiling point (bp) of the solventor co-solvent to maintain a Helmholtz double-layer duringelectrodeposition of a uniaxially oriented liquid crystal structure onthe surface of the semiconductor substrate.

During the sensor formation process, a second electrodeposition cycle isperformed to attach molecular recognition sites on top of the underlyinguniaxially oriented liquid crystal conducting polymer layer. The secondcycle is depicted in FIG. 4B. As with the first deposition cycle, adesired material is deposited from a medium; preferably a liquid medium422. In this case, second medium 422 contains headgroups 420, or anappropriate precursor, to be deposited. After medium 422 is brought intocontact with substrate 406 (to which polymer strands 410 were affixed inthe first cycle), a potential is applied to the substrate throughcircuit 408 to drive the second cycle. The potential will be positive ornegative depending upon the charge on the headgroups. This results indeposition of a lawn of immobilized headgroups 424 attached to anunfixed end of polymer strands 410. The potential and/or current densityshould be controlled to ensure that (1) the headgroup is affixed to thepolymer strands with strength to allow electron transport, and (2) thedeposited headgroups have a substantially common orientation. Depositioncycle variables are adjusted to ensure production of a single molecularlayer of uniaxially oriented liquid crystal chemically or biologicallyactive molecular recognition sites 424 individually “wired” tounderlying uniaxially oriented liquid crystal electrically conductingpolymer layer 410. The headgroups to be deposited may be provided withone or more functional groups which direct the headgroups onto strands410 in a desired orientation. As with the polymer strands, theheadgroups may be functionalized with a charge group. In many cases, itmay be desirable to locate the charge group away from the active site ofthe headgroup, so that the headgroup will attach with the active sitefacing the medium.

Deposition conditions must be tailored to the material to be deposited.In one embodiment of this invention, DNA deposition and GOD enzymedeposition conditions happened to use similar current density andapplied potential (e.g., 10 to 300 μA cm⁻² and 10 to 300 mV). However,solution conditions in the two deposition cycles (i.e., concentration ofdeposited material, pH, electrolyte, solvent) will not be the same.

As should be apparent, the deposition reactions require that the polymerstrands and recognition headgroups be electrically charged and mobile inan electric field. Thus, the compositions of the first and second mediamay have to be carefully chosen. Typically, though not necessarily, thefirst medium is removed and the substrate is allowed to dry before beingcontacted with the second medium.

1. Deposition of Uniaxially Oriented Liquid Crystal ConductingBiopolymers (Proteins and DNA)

In a preferred embodiment of the present invention, EMOLE methods areused to sequentially deposit, attach, and orient liquid crystalconducting polymers (e.g., DNA and proteins) onto the surface of asubstrate (e.g., a p-type silicon of a polycrystalline p-n junctionsolar cell). For example, the pH of a DNA-electrolyte depositionsolution is adjusted to ˜6.0 (more than two pH units above the pK_(a) orpI of DNA) producing a negatively charged DNA biopolymer. Light inducedphotoconduction by a submerged solar cell generates an electric field inthe DNA-electrolyte solution which uniaxially orients negatively chargedDNA strands onto the positive p-type silicon surface. Solar cell appliedcurrent density and voltage potential are low enough to establish andmaintain a Helmholtz double-layer (as described in U.S. Pat. No.5,215,631 to Westfall) between the p-type silicon surface and the DNAand counter ions in solution. The very gentle EMOLE conditionsfacilitate electrochemical deposition of uniaxially oriented liquidcrystalline extended DNA structures orthogonal to the semiconductorsubstrate surface. By “gentle,” it is meant that the conditions preservethe Helmholtz double-layer as described in the Westfall referencediscussed above.

EMOLE methods may be used to sequentially deposit, attach, and orientliquid crystal conducting protein (i.e., molecular recognition sites) ontop of the underlying uniaxially oriented liquid crystal DNA layeraffixed to the surface of the silicon substrate chip. For example, thepH of a protein-electrolyte deposition solution is ˜7.0 (more than twopH units above the pK_(a) or pI of the protein) producing a negativelycharged protein biopolymer. Light induced photoconduction by a submergedsolar cell generates an electric field in the protein-electrolytesolution which uniaxially orients negatively charged proteins onto the“lawn” of liquid crystal DNA molecular wires. Solar cell applied currentdensity and voltage potential are low enough to establish and maintain aHelmholtz double-layer between the DNA-modified p-type silicon surfaceand the protein and counter ions in solution. The very gentle EMOLEconditions facilitate sequential electrochemical depositions thatmaintain the first monolayer of uniaxially oriented liquid crystallineextended DNA structures orthogonal to the semiconductor substratesurface while depositing a second monolayer of uniaxially orientedliquid crystalline protein “headgroups” on top of the underlying “lawn”of liquid crystal DNA wires as characterized by the followingreferences: Collings, P J: Chap. 3. Electric and Magnetic Field Effects.In: Liquid Crystals: Nature's Delicate Phase of Matter. PrincetonUniversity Press; Princeton, N.J.; 1990; pp. 35-55. Collings, P J: Chap.9. Polymer Liquid Crystals. In: Liquid Crystals: Nature's Delicate Phaseof Matter. Princeton University Press; Princeton, N.J.; 1990; pp.162-180. Pelzl, G: Chap. 2. Thermodynamic Behavior and PhysicalProperties of Thermotropic Liquid Crystals. In: Liquid Crystals.Stegemeyer, H; guest ed. Steinkopff, Darmstadt and Springer, New York;1994; pp. 51-102. Zentel, R: Chap. 3. Liquid Crystalline Polymers. In:Liquid Crystals. Stegemeyer, H; guest ed. Steinkopff, Darmstadt andSpringer, New York; 1994; pp. 103-141).

Upon electrochemical deposition of a monolayer of uniaxially orientedliquid crystal protein, the DNA-silicon substrate is removed from thedeposition bath and allowed to slowly dry and cool in the presence of anapplied electric field. This allows the oriented liquid crystal proteinstructure to be “locked-in” on top of the oriented liquid crystal DNAmolecular wire terminal surface of the dry silicon substrate chip asdescribed in the following references: Collings, P J: Chap. 6. LiquidCrystal Displays. In: Liquid Crystals: Nature's Delicate Phase ofMatter. Princeton University Press; Princeton, N.J.; 1990; pp. 96-120.Albrecht, C; Enkelmann, V; Lieser, G; Schwiegk, S; Wang, W; Wegner, G;Zierer, D: The Crystallization Behavior of Rod-Like Macromolecules. In:Crystallization of Polymers. Dosiere, M; ed. Kluwer Academic Publishers;Dordrecht, Boston, London; 1993; pp. 323-330. Brandes, R: Part I.Generation of Tailored Radio Frequency Pulses For NMR. Part II.Deuterium NMR Studies of Oriented DNA, and Its Interaction With Water.Dissertation, Ph.D. in Chemistry; University of California, San Diego;1988.

Because EMOLE employs an electrodeposition mechanism, the species to bedeposited must be charged. Such charge exists naturally on manymaterials of interest when in the solution phase. However, manymaterials must be charged to facilitate EMOLE deposition. Manybiopolymers, for example, can be positively charged by adjusting the pHof the biopolymer-electrolyte deposition solution to more than two pHunits below the pK_(a) or pI of the biopolymer. The resulting positivelycharged species is suitable for electrochemical deposition onto negativen-type semiconductor surfaces, for example.

Like all liquid crystals, the oriented polymers of this invention mayhave their properties tailored by adding suitably functionalized groupsof atoms to the polymer backbone. Such properties include mechanicalstrength as well as ferroelectricity, non-linear optical activity, andelectronic charge transfer. The physical principles involved aresummarized in a number of books (Collings, P J: Liquid Crystals.Nature's Delicate Phase Of Matter. Princeton University Press;Princeton, N.J.; 1990. Stegemeyer, H (guest ed.): Liquid Crystals.Steinkopff, Darmstadt and Springer, New York; 1994. Plate, N A (ed.):Liquid-Crystal Polymers. Plenum Press; New York, London; 1993. Dosiere,M (ed.): Crystallization of Polymers. Kluwer Academic Publishers;Dordrecht, Boston, London; 1993). Anisotropic chemical and physicalproperties of liquid crystals and liquid crystal polymers are a resultof the molecular-scale structure formed. It was recently realized thatmanipulation of molecular-scale structure, and therefore function ofliquid crystals and liquid crystal polymers, not only depended on theuse of different functionalized organic molecules, but was heavilydependent on variables such as solvent, electrolytes, impurities,dopants; liquid crystal field effects (i.e., applied electric, magnetic,temperature, mechanical, electromagnetic radiation, or chemical fields);and processing techniques used (Collings, P J: Liquid Crystals. Nature'sDelicate Phase Of Matter. Princeton University Press; Princeton, N.J.;1990. Stegemeyer, H (guest ed.): Liquid Crystals. Steinkopff, Darmstadtand Springer, New York; 1994. Plate, N A (ed.): Liquid-Crystal Polymers.Plenum Press; New York, London; 1993. Dosiere, M (ed.): Crystallizationof Polymers. Kluwer Academic Publishers; Dordrecht, Boston, London;1993. Collyer, A A (ed.): Liquid Crystal Polymers: From Structures ToApplications. Elsevier Applied Science; London, New York; 1992. Lam, L;Prost, J (eds.): Solitons In Liquid Crystals. Springer-Verlag; New York,Berlin, Heidelberg, London; 1992). For example, coupling of molecularrecognition surfaces to electronically conducting polymers may resultfrom chiral smectic (layered cholesteric) liquid crystal structuresformed by sequential deposition of DNA and protein using EMOLEfabrication techniques provided by this invention. In a preferredembodiment, biopolymers (DNA and protein) and EMOLE techniques are usedto fabricate a molecular recognition (MR) device.

IV. CONDUCTING POLYMERS AND THIN FILMS

Many different conducting polymers and thin films can be employed for“wiring” molecular recognition sites to a semiconductor or standardelectrical component substrate. Generally such polymers may bebiological, organic, inorganic, water soluble, lipid soluble orcombinations thereof. Many examples of conducting polymers suitable forthis invention are discussed by Skotheim, T A: Handbook Of ConductingPolymers. Vol. 1-2. Skotheim, T A, ed. Marcel Dekker, Inc.; New York,Basel; 1986. Types of conducting polymers and thin films suitable foruse in this invention include, but are in no way limited to thefollowing general classes: aromatic metal-doped polymers (e.g.,polyaniline doped by metal salts), π-stacked (aromatic) polymers (e.g.,polyphenanthroline; pyrazine-bridged polymers of π-stackedmetalloporphyrins; 2,3,6,7,10,11-hexahexylthiotriphenylene (HHTT)),π-stacked (aromatic) helical polymers (e.g., DNA), organic π-conjugatedlinear polymers (e.g., polyacetylene), heterocyclic polymers (e.g. DNA,polyporphyrins), macrocyclic polymers (e.g., polyporphyrins with a redoxmetal; polytetrazacyclododecane with a redox metal), porphyrin polymers,polymer composites (e.g., layered polymer mixtures), polyelectrolytepolymers, (e.g., proteins, DNA), liquid-crystal polymers (e.g., certainproteins; DNA; polyporphyrins; and2,3,6,7,10,11-hexahexylthiotriphenylene (HHTT)), self-organizingpolymers (e.g., polysurfactants with redox metal; HHTT), branchedpolymers, dendritic polymers (e.g., starburst dendrimers with redoxmetal), chaotic polymers (e.g., poly (SiO₂)_(n) in glass with redoxpolymer), biopolymers (e.g., protein, DNA, polyporphyrins), inorganicpolymers (e.g. iron (hydrous) oxides), organometallic polymers (e.g.,ferrocene polymers), inorganic/organic hybrid polymers (e.g. iron(hydrous) oxide/polybipyridine complex), metallocene polymers (e.g.,polyferrocene), inclusion compound polymers (e.g., polyzeolite withredox metal), mixed doped polymers, colloidal/sol-gel doped polymers(e.g., poly (SiO₂)_(n) with redox metal), ionomers (e.g., DNA, certainproteins, and certain polysurfactants), metal cluster doped polymers(e.g., iron (hydrous) oxide/polybipyridine complex), redox polymers(e.g., Heller (Osmium-PVP) and Skothiem (ferrocene-polysiloxane)), blockpolymers, graft polymers, transition metal films (e.g., deposited byatomic layer epitaxy (ALE)), high temperature superconductor films(e.g., atomic layer epitaxy (ALE) of appropriate redox metals),Langmuir-Blodgett films (e.g., detergents, amphiphiles, surfactants),sol-gel glass films (e.g., spin glass films), etc., or any combinationsof the above. Conducting polymers of appropriate strand lengths for eachof these may be employed herein.

In some cases, the native form of the polymer will be an insulator, butupon appropriate doping, addition of impurities, hydration,conformational change, ionization, oxidation, reduction, etc. theybecome conductive. Further, some conducting polymers may be reversiblyswitched between conductive and insulative states. Polyaniline, forexample, will become conductive in the protonated or oxidized form.Other “switchable” conductive polymers include, for example, polymerspolymerized from the following monomers: N-methylpyrrole, thiophene,3-methylthiophene, 3,4-dimethylthiophene, vinylferrocene, styrene,nitrostyrene, viologens, vinyl-pyridine, vinyl-2,2′-bipyridine,vinylrubrene, quinone-based compounds, and derivatives thereof. Thisinvention may also take advantage of such conductivity transformation asa primary or auxiliary sensing mechanism. For example, a sensor signalmay only be triggered by a combination of two events: a ligand bindingwith a molecular recognition headgroup and a pH change which causes thepolymer wiring to become conductive.

Enzymes used in organic synthesis (i.e., to produce drugs andpharmaceuticals), may be used as molecular recognition headgroups ofthis invention. These include, but are not limited to, combinatorial andcommercial libraries of esterases, lipases, amidases, acylases, andother thermophillic and mesophilic enzymes with broad substratespecificities that can catalyze reactions in organic solvents and athigh temperatures. Upon ligand binding to an esterase or lipase, areaction will take place producing an alcohol and carboxylic acid fromthe cleaved ester bond. This will make the pH of theheadgroup/switchable polymer molecular environment more acidic; thus,protonating a reversibly switchable polymer to the protonated orconducting form. Amidase or acylase cleavage of an amide bond willproduce a free amine and a carboxylic acid. Chelation of the acid byanion exchange support would leave an increasing concentration of freeamine which would make the pH of the headgroup/switchable polymermolecular environment more basic; thus, deprotonating the reversiblyswitchable polymer to the neutral or insulating form. Examples ofenzymes used in organic syntheses may be used as molecular recognitionheadgroups to monitor levels of drugs and pharmaceuticals in the humanblood.

Esterase, lipases, acylases, or amidases may also be used to deprotectligands to alcohols, carboxylic acids, or free amines which then becomesubstrates suitable for a second molecular recognition headgroup, usedto produce a signal by methods described in the present invention. Forexample, cholesterol esterase cleaves cholesterol ester found in bloodto cholesterol, which is then a substrate for cholesterol oxidase.Cholesterol oxidase would produce a signal much like glucose oxidasedescribed as an example of this invention.

Other approaches include, for example, Swager, et al. (Swager, T M;Marsella, M J; Conducting Polymers With Chemical Sensitive Traps andBarriers: New Molecule-Based Sensors. Mat. Res. Soc. Symp. Proc.328:263-266, 1994) which describes reversibly switchable polythiophenederivatives which exhibit large changes in bandgap in the presence ofspecific ions. These materials are based upon novel crown etherscontaining bithiophene monomers. Sensory polymers which are selectivefor K⁺ and Na⁺ are described. In such materials, specific ions induce atwisting of the polymers backbone, resulting in a decrease of 1-orbitaloverlap between thiophene rings; reducing the extent of conjugationgiving rise to an insulating (higher bandgap) form.

Another example is of a sequence-specific DNA sensor. A specificsequence of single-strand DNA (nonconducting or insulating form) with 5′or 3′ terminus thiol could be adsorbed to a gold electrode substrate. Ananalyte sample containing the complementary DNA sequence would produce aDNA double-strand polymer which is a conducting form of DNA. This resultis a DNA sequence detector. DNA of the wrong sequence would not produceDNA double-strand polymer (conducting form). Appropriate end groupfunctionalities on single-strand DNA or no end group modifications ofsingle-stranded DNA (i.e., native DNA) using EMOLE methods could be usedto put sequence-specific single-strand (insulating form) DNA onsemiconductor substrates for use as a DNA sequence detector. DNA atcrime scenes could be identified on the spot, doing away with PCRtechniques and laborious and very costly DNA sequencing laboratoryprocedures.

Chemical-, photo-, or electro-polymerization of monomers may take placedirectly on the semiconductor or standard electrical component substratesurface or pre-polymerized polymers may be deposited. Furthermore, onceattached and polymerized, the polymer or thin film may be oriented intoa highly conductive liquid crystal polymer or thin film form. This maybe accomplished by depositing polymers in the presence of appropriateelectrical, magnetic, or chemical (solvent) fields. Preprocessing orconditioning of polymers is described in the Handbook of PolymerSynthesis (Plastics Engineering Series, Volume 24) Kricheldorf, H. F.,1991. Chemical polymerization may employ, for example, H₂O₂,organoperoxides, or 2,2′-azobisisobutyronitrile (AIBN).Photopolymerization may employ photons which generate photochemicalradicals which can initiate and propagate polymerization.Electropolymerization is currently employed to synthesize conductingpolymers.

A. Electron Transport Proteins

An example of a conducting biopolymer that may be useful for thisinvention is the electron transport protein. Electron transport proteinsare a product of millions of years of biological evolution, fine tuningthe function of electronic conduction. In nature, electron transportproteins often reside in, and are oriented by, a liquid crystallinelipid bilayer membrane. In this invention, the electron transportprotein may be deposited into a close-packed oriented two-dimensionalcrystalline structure by EMOLE crystallization processing techniques.This produces a surface structure suitably oriented as a plurality ofmolecular wire interconnects.

Proper deposition and orientation of proteins can be accomplished bymanipulation of the physical and chemical conditions duringcrystallization. The EMOLE technique allows a systematic approachunderstanding and optimizing the relevant parameters for depositingprotein or peptide polymers as wires for sensors. More generally, thenewly developed techniques of EMOLE provide for experimental control ofprotein crystal structure and function.

Electron transport proteins are in some embodiments suitable for usewith this invention because they perform some of the function desiredfor molecular electronic device (MED) fabrication—i.e., electron storageand transfer at the molecular-scale. These properties arise from thealpha-helical and beta-pleated sheet structures of these biologicalmacromolecules and from their non-protein prosthetic groups. Theseprosthetic groups are inorganic-, organometallic-, or metal atomcofactors which are integral to the structure of protein. A particularlyinteresting protein is cytochrome b₅₆₂ of E. coli. This protein is small(12,000 daltons), has a single polypeptide chain folded into a simple4-alpha-helical motif, the x-ray structure is known to 2.5 Å, and mostimportantly, the single heme group is non-covalently bound. This lastproperty allows for the substitution of other porphyrin analogs with avariety of coordinated metal atoms, greatly increasing the experimentalflexibility of the system (Ulmer, K M: Chap. 29. Self-Organizing ProteinMonolayers As Substrates For Molecular Device Fabrication. In: MolecularElectronic Devices II. Carter, F L; ed. Marcel Dekker, Inc.; New York,Basel; 1987; pp. 573-590).

Photosynthetic electron transport proteins electronically connectingphotosystem II and photosystem I in plants, and mitochondrialrespiratory electron transport proteins are examples of conductingbiopolymer proteins oriented by a liquid crystalline lipid bilayermembrane—the chloroplast membrane (Clayton, R K: Light and LivingMatter, Volume 2: The Biological Part. McGraw-Hill Book Company, NewYork, 1971) and mitochondrial membrane; facilitating an extremelyefficient electron transfer chain via electron tunneling mechanism(Pethig, R: Chap. 9. Electronic Properties of Biomacromolecules. In:Dielectric and Electronic Properties of Biological Materials. John Wiley& Sons; Chichester, New York; 1979; pp. 290-356).

Electron transport proteins that may be found among the proteinsparticipating in the respiratory chain of mitochondria are for example:flavoproteins, nonheme iron proteins, and cytochromes b, c₁, c, a, anda₃. With the exception of the electron donor, NADH, all of these areelectron transport proteins, shuttling two electrons from each moleculeof NADH to reduce ½O₂ to H₂O. This downstream free energy electrontransport to O₂ is coupled to phosphorylative production of ATP, abiochemical energy currency.

Electron transport from photosystem II to photosystem I in thechloroplast membrane of green plants involves the electron transportproteins cytochrome b₅₅₉ or b₃ and cytochrome f. Electron transport fromphotosystem I involves the electron transport proteins ferredoxin andcytochrome b₆.

All of these electron transport proteins are juxtaposed to each other inmembranes with increasing standard oxidation-reduction potentialsfacilitating a downward free energy transfer of two electrons from oneelectron transporting protein to the next in a highly ordered chain.

B. DNA Quantum Wires

A second example of a conducting biopolymer not normally thought of aselectrically conductive until recently is DNA (Meade, T J and Kayyem, JF: Electron Transfer Through DNA: Site-Specific Modification of DuplexDNA with Ruthenium Donors and Acceptors. Angew. Chem. Int. Ed. Engl.34(3):352-354, 1995. Murphy, C J; Arkin, M R; Jenkins, Y; Ghatlia, N D;Bossmann, S H; Turro, N J; Barton, J K: Long-Range Photoinduced ElectronTransfer Through a DNA Helix. Science 262:1025-1029, 1993. Meade, T J:Chap. 13. Electron Transfer Reactions Through the DNA Double Helix. In:Metal Ions In Biological Systems. Vol. 32. Interactions of Metal IonsWith Nucleotides, Nucleic Acids, and Their Constituents. Sigel, A;Sigel, H; eds. Marcel Dekker, Inc.; New York, Basel, Hong Kong; 1996;pp. 453-478. Stemp, EDA; Barton, J K: Chap. 11. Electron TransferBetween Metal Complexes Bound To DNA: Is DNA A Wire? In: Metal Ions InBiological Systems. Vol. 33. Probing of Nucleic Acids by Metal IonComplexes of Small Molecules. Sigel, A; Sigel, H; eds. Marcel Dekker,Inc.; New York, Basel, Hong Kong; 1996; pp. 325-365. Arkin, M R; Stemp,EDA; Holmlin, R E; Barton, J K; Hormann, A; Olson, EJC; Barbara, P F:Rates of DNA-Mediated Electron Transfer Between Metallointercalators.Science 273:475-480, 1996). DNA is a biopolymer with known solution andsolid-crystal structures. In this invention, deposition of an orientedextended liquid crystalline DNA structure orthogonal to asolid-substrate surface may be achieved by EMOLE crystallizationprocessing techniques. This produces a surface structure suitablyoriented as a plurality of molecular wire interconnects.

While not wishing to be bound by theory, the following discussion ispresented to illustrate the state of the art as to DNA as a conductingmedium. There is still no consensus in the art as to whether DNA canactually act as a wire. The debate is set forth generally by Wilson(Wilson, DNA: Insulator or Wire, Chem. & Eng. News, 1997:33, Feb. 24,1997). While such debate rages, the following discussion assumes thatDNA is in fact a very good conducting polymer and is a preferred wirefor use with the sensors and EMOLE methods of this invention.

Long distance electron movement through DNA (i.e., ˜40 Å or ˜12 basepairs) has been confirmed only in experiments in a water solution. DNAhas to be fixed to a terminal base, substrate, etc., coupled with thecontrolling of the thickness and orientation of molecules in order tomeasure the accurate conductivity of the fixed DNA. Recently, studies onfixation of DNA to solid bases have been reported by various methodssuch as ion connection, covalent bond, and protein bonding for the useof DNA as a potential electronic material.

Okahata, et al. prepared a polyion complex using DNA and cation lipidsin order to prepare thin cast film membranes of DNA (Ijiro, K andOkahata, Y: A DNA-Lipid Complex Soluble in Organic Solvents. J. Chem.Soc., Chem. Commun. 1992:1339, 1992). Phosphate and cation lipids formedquantum chemical ion pairs. As a result, an alkyl base covered the DNAforming the shape of a brush to wash a test tube and became hydrophobicand settles instantly. Nishi et al. prepared the gel film with thethickness of 2-3 μm×2-3 mm by adding bivalent metallic ions such as Ca²⁺or Mg²⁺ to a water solution of alginic acid, a polysaccharide having aresidue of alginic acid (Iwata, K; Nishi, N; Miura, Y; Nishimura, S;Tokura, S: Polymer Preprints, 42:599, 1993). DNA structure wasmaintained in the film from adsorption test of intercalator color in thestudy. The molecular orientation of DNA in the film prepared by fixationmethods was random and was very difficult to control the molecularorientation and thickness of the membrane. G. Decher et al. reported onthe methods for preparing the thin membrane of DNA which had a thicknessof one molecule (Lvov, Y; Decher, G; Sukhorukov, G: Assembly of ThinFilms by Means of Successive Deposition of Alternate Layers of DNA andPoly(Allylamine). Macromolecules 26:5396-5399, 1993). High molecularweight DNA isolated from sturgeon sperm formed layers 33 Å thick byx-ray diffraction indicating the DNA spread two-dimensionally with thelong axis parallel to the substrate surface. In conventional studies,fixation was performed using the ion connection of anion phosphates atmultiple points. On the other hand, Maeda et al. reported the fixationmethods fixed the special edge of DNA on a gold terminal by chemicallytreating the edge of DNA with a thiol base (Maeda, M; Nakano, K; Uchida,S; Takagi, M: Mg²⁺-Selective Electrode Comprising Double-Helical DNA asReceptive Entity. Chem. Lett. 1994:1805-1808, 1994). Organic thiolcompounds bind strongly to gold. Maeda et al. considered that theorientation of DNA was vertical towards the terminal from themeasurement of the amount of fixed DNA. Ijiro et al. reported aproduction of a semi-molecular membrane using DNA, a cation intercalatorlipid (C₁₈-acridine orange), and Langmuir-Blodgett techniques of castinga thin film. Orientation of the DNA strings was attempted by applyingcompression and measuring conductivities in different directions (Ijiro,K; Shimomura, M; Tanaka, M; Nakamura, H; Hasebe, K: Thin Solid Films (inpress). Ijiro, K and Shimomura, M: Double-Stranded DNA for MolecularElectronic Devices. Kotai Butsuri 30(12):1042-1048, 1995. Birdi, K S:Lipid and Biopolymer Monolayers at Liquid Interfaces, Plenum Press; NewYork, London; 1989). As evidenced by this review of various methods forfixation of DNA on surfaces, there is some difficulty in orienting DNAfilms for use as routine commercial electronic materials providing highdensity molecular wire interconnects on common semiconductor or standardelectrical component substrates.

In a preferred embodiment of this invention, DNA or nucleic acid is usedas the conducting polymer precursor to be electrochemically depositedand uniaxially oriented into a highly conductive liquid crystalline formon the semiconductor substrate surface. Single-stranded DNA is notelectrically conductive as a molecular wire. It is a random coil withlittle order. However, double-stranded A-, B-, or Z-DNA are examples offlat heteroaromatic purine and pyrimidine π-stacked base pairs (i.e.,heteroaromatic r-stacking of flat base pairs, one on top of the next ina rising helix) that makes double-stranded DNA conductive. Otherexamples of suitable DNA structures that may be deposited as uniaxiallyoriented liquid crystalline DNA quantum wires include, but are in no waylimited to clockwise double-stranded twining structures, otherwisecalled A-, B-, C-, D-, E-, and T-types. DNA also has a counterclockwisedouble-stranded twining structure, called Z-type. In addition, there islooped DNA which consists of thousands of pairs of bases called plasmidDNA which exists in prokaryotic organisms. There is also a twistedlooped DNA structure which comprises several loops and a super helicalstructure. There even exists a twisted loop, cross shaped DNA (Ijiro, Kand Shimomura, M: Double-Stranded DNA for Molecular Electronic Devices.Kotai Butsuri 30(12):1042-1048, 1995). And DNA exists in triple helixtype structures as well (Povsic, T J; Dervan, P B: Triple HelixFormation By Oligonucleotides On DNA Extended To The Physiological pHRange. J. Am. Chem. Soc. 111(8):3059-3061, 1989).

Preferably, a liquid crystal B-DNA type double-stranded structure isdeposited, electrically attached, and uniaxially oriented in parallelextended conformation orthogonal to the surface of a semiconductor inspecific chemically or electrochemically activated regions (as shown inFIG. 2). A and T; G and C complementary pairs of bases form an uprightduplex helical structure with a diameter of approximately 20 Å,comprising two high molecular chains. The pitch of the duplex helicalstructure is approximately 34 Å and 10 of the pairs of bases line upvertically towards the extended line of DNA. The upper and lower pairsof bases create an angle of 36° while the distance between each pair ofbases is 3.4 Å. This produces a strong mutual relationship between eachstuck pair of bases inside the duplex helical structure of DNA. Forexample, an extreme reduction of absorbance (light color effect) willoccur because of π-π* conversion. In other words, the internalcharacteristics of DNA can be considered as a suspected one-dimensionalcrystalline structure of stuck pairs of bases (Ijiro, K and Shimomura,M: Double-Stranded DNA for Molecular Electronic Devices. Kotai Butsuri30(12): 1042-1048, 1995).

Particularly high packing efficiencies are achieved in the icosahedraldouble-stranded DNA bacteriophages, where the DNA duplexes are closepacked at a center-to-center spacing of about ˜26 Å. This constraint hasbeen incorporated into several recent models in all of which the rods ofduplex DNA are configured in more-or-less parallel bundles (Booy, F P;Newcomb, W W; Trus, B L; Brown, J C; Baker, T S; Steven, A C:Liquid-Crystalline, Phage-Like Packing Of Encapsidated DNA In HerpesSimplex Virus. Cell 64:1007-1015, 1991). Moreover, the average 26 Åinterduplex spacing closely resembles that observed for liquid crystalsof DNA in vitro by cryoelectron microscopy or x-ray diffraction (Booy, FP; Newcomb, W W; Trus, B L; Brown, J C; Baker, T S; Steven, A C:Liquid-Crystalline, Phage-Like Packing Of Encapsidated DNA In HerpesSimplex Virus. Cell 64:1007-1015, 1991). In a preferred embodiment ofthis invention, uniaxially oriented liquid crystalline B-DNA conductivewires are electrochemically deposited at specific light activatedregions on the surface of a p-n junction solar cell by EMOLE fabricationmethods as described above.

V. MOLECULAR RECOGNITION SURFACES

A molecular recognition surface preferably is made up of atwo-dimensional crystal array of one or more molecular recognitionsite(s) that recognize a particular ligand (i.e., analyte) typically,though not necessarily, in a liquid. In addition to its ability to bindspecific ligands, a molecular recognition site may also be a catalyticsite, redox site, electron transfer site, energy transfer site, magnetictransfer site, and as a consequence of ligand binding may induceconformational change, and quantum-confined electron/hole tunneling andpercolation.

The molecular headgroups employed in this invention include, forexample, proteins (which bind ligands), catalytic antibodies,porphyrins, lectins, enzymes (including any enzyme categorized in the ECNomenclature—e.g., class 1: oxidoreductases, class 2: transferases,class 3: hydrolases, class 4: lyases, class 5: isomerases, and class 6:ligases), immunological antibodies, antigens, receptors, viruses, cells,cavitands, zeolites (which bind redox metals), supramolecularassemblies, electro-optical materials (e.g., second- and third-ordernonlinear optical materials), photoconductive and photoelectricmaterials (in which an applied electromagnetic field produces freeelectrons), giant magnetoresistive materials (in which an appliedmagnetic field changes resistivity of the material), metal chelates,magnetic materials (in which magnetic ordering is changed by thepresence of other magnetic materials), inorganic scintillators (whichconvert high energy radiation to lower energy light photons), inorganiccrystal oscillators (which act as a quantum frequency transmitter andreceiver), piezoelectric materials (in which mechanical force produceselectron flow), light-harvesting polymer systems (in which lightproduces electron flow and chemical energy storage), laser switch dyes(which absorb light at one wavelength and emit a monochromatic light ata longer wavelength), barrier tunnel switches (e.g., molecular electronswitches), etc.

Examples of ligands that can be used with this invention include, butare not restricted to, agonists and antagonists for cell membranereceptors, toxins and venoms, viral epitopes, antigenic determinants,monoclonal and polyclonal antibodies, hormones, hormone receptors,steroids, peptides, enzymes, substrates, cofactors, drugs, lectins,sugars, oligonucleotides, oligosaccharides, proteins, transition metals,chelates, cavitands, pollutants, chemical and biological warfare agents,poisons, dyes, gases, intercalators, alcohols, alkaloids, fats, lipids,cholesterol, blood type, cell surfaces, metabolites, etc.

Molecular recognition sites that mediate a biological or chemicalfunction either directly or indirectly on binding with a particularligand(s) are of most interest. Suitable molecular recognition sitesinclude relatively small, single molecules, such as cofactors, whichshow specific binding properties. Typically, molecular recognition siteswill range from 1 dalton to greater in size. Other examples of molecularrecognition sites include, but are not restricted to, the common classof receptors associated with the surface membrane of cells and include,for instance, the immunologically important receptors of B-cells,T-cells, macrophages and the like. Other examples of molecularrecognition sites that can be investigated by this invention include butare not restricted to hormone receptors, hormones, drugs, cellularreceptors, membrane transport proteins, electron transport proteins,steroids, peptides, enzymes, substrates, cofactors, vitamins, lectins,sugars, oligonucleotides, intercalators, oligosaccharides, viralepitopes, antigenic determinants, glycoproteins, glycolypoproteins,immunoglobins, restriction enzymes, catalytic antibodies, transitionmetals, chelates, cryptands, cavitands, supramolecular structures, etc.

A. Oxidoreductases (Redox Enzymes)

Examples of molecular recognition sites that bind specific ligands,catalyze a redox reaction, and are electrically conducting biopolymers,are a broad class of enzymes called the oxidoreductases. To this classbelong all enzymes catalyzing oxido-reductions. The substrate oxidizedis regarded as hydrogen or electron donor. The classification is basedon ‘donor:acceptor oxidoreductase’. The recommended name is‘dehydrogenase’, wherever this is possible; as an alternative, ‘acceptorreductase’ can be used. ‘Oxidase’ is only used in cases where O₂ is anacceptor. Classification is difficult in some cases because of the lackof specificity towards the acceptor. The EC number 1.x.x.x as it appearsin Enzyme Nomenclature (1978) is assigned to the class calledoxidoreductases (Enzyme Nomenclature. Academic Press; New York; 1978).

Oxidoreductases or redox enzymes are molecules of 40,000 daltons (e.g.,galactose oxidase) to 850,000 daltons (e.g., choline dehydrogenase) withone or more redox centers. Their average hydrodynamic diameters rangefrom ˜55 to ˜150 Å. In the great majority of enzymes, the redox centersare located sufficiently far from the outermost surface (defined byprotruding protein or glycoprotein domains) to be electricallyinaccessible. Consequently, most enzymes do not exchange electrons withelectrodes on which they are adsorbed, i.e., their redox centers areneither electrooxidized at positive potentials nor electroreduced atnegative ones. Apparently, part of the protein or glycoprotein shellsurrounding the redox centers is there to prevent indiscriminateelectron exchange between the different redox macromolecules of livingsystems. Another function of this shell is to stabilize the structure ofthe enzyme. Because neither function is essential for catalysis, redoxenzymes do function when part of the shell is stripped or, when theshell is chemically altered so as to make it electrically conductive.

Examples of oxidoreductase enzymes suitable for use with this inventioninclude glucose oxidase, catalase, peroxidase, cholesterol oxidase, andalcohol dehydrogenase. Glucose oxidase (GOD) turns over at ambienttemperature at a rate of ˜10² s⁻¹, i.e., it produces about 200transferable electrons/s. Because it has a radius of ˜43 Å, there can beup to 1.7×10¹² enzyme molecules on the electrode surface. The currentdensity, when all redox centers are electrically well connected to theelectrode, may thus reach about 3.4×10¹⁴ electrons s⁻¹ cm⁻², or 53 uAcm⁻².

In a preferred embodiment, molecular recognition site(s) will becomposed of one or more of the following oxidoreductases (redoxenzymes): glucose oxidase (GOD) which binds specifically to D-glucose,cholesterol esterase/cholesterol oxidase (COD) which binds specificallyto cholesterol ester/cholesterol, catalase (CAT) which bindsspecifically to H₂O₂, or alcohol dehydrogenase (ADH) which bindsspecifically to ethanol. All of these redox enzymes oxidize theirrespective substrates, transferring two electrons to natural orartificial diffusible electron acceptor mediators. In the presentinvention, a uniaxially oriented liquid crystal conducting biopolymer inan extended straight conformation is stuck or “wired” into eachcatalytic site/redox center permitting direct electron transfer to takeplace. Electron transfer to natural diffusible electron acceptors suchas O₂ or other artificial diffusible redox mediators such as ferroceneor metal derivatives is therefore largely eliminated. Mechanism ofelectron transfer in the present invention is based on a solid-state“hard-wired” organization at the enzyme catalytic site/redox centerestablishing quantum-confined electron/hole tunneling and percolationthrough a uniaxially oriented liquid crystal conducting polymer orbiopolymer known as a molecular or quantum wire. Electron or holeinjection from a molecular recognition headgroup (i.e., oxidoreductase)through an attached superconducting quantum wire tail (i.e., DNA)interconnect to an underlying electronic substrate is the basis of amolecular transistor.

In a preferred embodiment of this invention, a plurality of suchmolecular recognition sites (i.e., enzymes) are electrochemicallydeposited onto the surface of a p-n junction solar cell by firstdepositing liquid crystalline highly oriented B-DNA “molecular wires” tothe p-type surface. Preferably, a liquid crystalline molecularrecognition surface structure is deposited, electrically attached, anduniaxially oriented at the surface of a liquid crystal B-DNAdouble-stranded structure which was deposited, electrically attached,and uniaxially oriented at the surface of p-type semiconductor inspecific chemically or electrochemically activated regions. Oriented DNAduplex polyelectrolytes likely are extended, straight quantum wires thatpenetrate deeply into enzyme crevices at one end and semiconductorsubstrate at the other end. This type of molecular-scale structurelikely facilitates direct, quantum mechanical electron transfer betweenenzyme headgroups and semiconductor substrate.

In a preferred embodiment, spatially addressable electrochemicalactivation at specific regions on the surface of a p-n junction solarcell is achieved by light masking or photolithographic techniques forthe purpose of electrodeposition at specified locations on the chip. Ina preferred embodiment of this invention, liquid crystalline highlyoriented molecular recognition surfaces are electrochemically depositedat specific light activated regions on the surface of a p-n junctionsolar cell by EMOLE methods as described above. Preferably, DNA wires onthe p-n junction solar cell are exposed to light at specific regions toform electrical contacts with liquid crystal oriented molecularrecognition sites by EMOLE methods. This is repeated at differentregions on the semiconductor surface to pattern complex digital organicintegrated circuits (IC) of “wired” molecular recognition sites. Thefabrication scheme described above constitutes preferable productionmethods of a molecular recognition chip (MRC).

B. Immunoglobulins

If we are looking for a more general method of incorporatingnon-biological molecules into molecularly organized materials, then theimmunoglobulins or antibody molecules offer many attractive advantages.Using currently available monoclonal antibody technology, it is nowpossible to generate a specific immunoglobulin molecule capable ofbinding to almost any compound of interest. In accordance with thepresent invention, one could engineer crystals of antibody complexes inwhich it was possible to control the arrangement and orientation of thecomplexed molecules at the molecular-scale. There has already been areport of successful application of Langmuir-Blodgett techniques toproduce two-dimensional crystals of antibody molecules which may be usedfor MED development.

Examples of molecular recognition sites that bind specific ligands,catalyze a redox reactions, undergo conformational change and areelectrically conducting biopolymers, are a broad class of proteinscalled immunoglobulins. Catalytic antibodies are man-madeimmunoglobulins that can be engineered to possess all of the abovechemical and physical properties and specificity for a particularligand. In a preferred embodiment, immunoglobulins or catalyticantibodies may be deposited as molecular recognition headgroups onto DNAquantum wires using EMOLE crystallization processing techniques asdescribed above to fabricate molecular recognition (MR) devices on amacro-solid substrate.

VI. CONDUCTION MECHANISMS THROUGH POLYMERS ON SOLID SUBSTRATES

A. Energy Bands in Uniaxially Oriented Liquid Crystal ConductingBiopolymers (Proteins and DNA) and Semiconductor Substrates

Since Szent-Gyorgyi's report that biopolymers can work likesemiconductors, many researchers have pursued research on electronmovement through proteins. The potential for long-range electronmovement within a protein coupled with double helix DNA wastheoretically calculated from the point of quantum chemistry. Becauseionic impurities are present in DNA, the methods used to prepare solidpellets varied depending on the experiments and thus, reportedconductivities have varied between 10⁻⁴ and 10⁻¹⁰ mho·m⁻¹. A quantummechanical-based model also offers a possible explanation for theanomalously rapid long-range (i.e., ˜40 Å) photoelectron transferrecently observed by Barton and Turro et al. for donor and acceptorspecies intercalated into a DNA double helix.

There is no possibility of intrinsic conductivity in periodic andaperiodic polypeptide chains due to their large fundamental energy gap.This conclusion may appear, at first glance, to be a stumbling block tothe electronic conduction in proteins. It should however be noted thatmany other materials, the glasses, oxides and amorphous semiconductors,also have energy gaps sufficiently large to make them poor conductorsbut this has not prevented consideration of them in electronic terms andthe establishment of a considerable body of experimental and theoreticalevidence for long range electron transfer in them.

Since the bands in the density of states (DOS) curves of aperiodicchains are very broad with a few small gaps, there is a possibility ofextrinsic conduction on doping with electron acceptors (p-doping) orwith electron donors (n-doping) in these chains. To decide about thenature of extrinsic conduction (whether Bloch-type conduction or chargetransport through hopping) one needs to investigate the localizationproperties of the wave functions belonging to the energy levels in theupper part of the valence band region or the lower part of theconduction band region (these are the regions of interest if a chargetransfer is to take place in vivo due to the interaction of proteinswith electron acceptors or donors or with DNA). The possibility of thistype of charge transfer has also been suggested by Szent-Gyorgyi.

Quantum mechanical models proposed to estimate energy bands andelectronic conduction in proteins and DNA can be influenced by a numbersof external factors which tend to reduce or eliminate the bandgap andbroaden the width of estimated valence and conduction bands. This yieldsbiopolymers with metallic-like conduction properties. Such externalfactors include impurities, dopants, applied electric fields, appliedmagnetic fields, illumination (hv), hydration with H₂O, solvent,pressure, conformational changes, orientation, pH, electrolytes, localsurface charges, and injection of electron or holes directly into theconduction or valence bands of the biopolymer. Injection of electronsinto protein conduction bands can come from COO⁻ groups on protein sidechains or at the carboxyl terminus, and from H₂O. Selective applicationof these external factors effects are used to engineer bandgap structureof proteins and DNA using EMOLE fabrication techniques to producedesired physical and chemical properties of superconducting, conducting,semiconducting, or insulative forms. EMOLE provides energy band matchingand molecular interconnects between proteins, DNA, and the semiconductorsubstrate which affords quantum mechanical electronic conduction.

Uniaxially oriented liquid crystalline forms of conducting biopolymers(proteins and DNA) may be produced by EMOLE fabrication techniques.Processing variables utilized by EMOLE to deposit oriented liquidcrystal conducting biopolymers include external factors influencingbiopolymer energy band structures described above. EMOLE is a chipfabrication method used to engineer molecular structure, energy bandstructure, band matching, and quantum mechanical molecular interconnectsof conducting biopolymers (proteins and DNA) on the surface of asemiconductor substrate.

In some embodiments, a molecular plastic insulator is formed aroundconducting polymers and biopolymers (e.g., proteins and DNA) asdescribed herein and which may be produced by, for example, EMOLEfabrication techniques. The insulator may be separately provided at anylevel of resolution; in some embodiments it is applied separately atseparately addressable regions of sensor. Thus, depending on theapplication, the insulator may be applied in a manner that isolatesdistinct regions of the sensor at specific macro-, micro-, and/ornanometer scale. In such embodiments, as with other embodimentsdescribed herein, the conducting polymers and/or biopolymers may benon-randomly oriented (e.g., uniaxially oriented and even liquidcrystalline in some instances).

In order for communication between uniaxially oriented liquid crystalconducting biopolymers (proteins and DNA) and the polycrystalline ormonocrystalline macro-semiconductor substrate of the MR-device, commonenergy levels must exist between not only the protein (molecularheadgroup) and DNA (quantum wire tail) components, but between the DNAand the semiconductor substrate. In MR-devices of the present invention,DNA duplex polyelectrolytes are extended, straight quantum wires thatpenetrate deeply into enzyme crevices at one end, and into themacro-semiconductor substrate at the other end. This type ofmolecular-scale structure facilitates direct electron transfer from theenzyme prosthetic group and desired energy continua between enzyme, DNA,and semiconductor substrate. The nature of the energy continua issimilar to the ideas proposed by Szent-Gyorgyi in 1946, Pethig,Bathaski, Tanatar, and others regarding a common quantum mechanicalenergy band continuum, resonant tunneling, hopping, acoustic plasmon,etc. mechanisms which facilitate charge transfer of mobile chargecarriers (electrons or holes) through protein and DNA to the underlyingsemiconductor substrate.

B. Superconductivity

The possibility that superconductive phenomena may play a biologicalrole is at present a controversial subject in several laboratories.Unlike the situation for normal electronic conductors, electrons in asuperconductor are not free to move independently of each other butexist as coupled electron pairs constrained to be in the same quantumstate. As a result of this pairing-up of electrons, electron scatteringeffects are minimized with the result that the flow of electron currentcan occur without the generation of heat and hence with no electricalresistance. Such an effect could obviously have far-reachingconsequences if it could be detected in biological systems atphysiological temperatures. In conventional superconductors the electronpairing results from interactions between the electrons and the latticephonons. In 1964, Little proposed that suitably constructed organicpolymeric systems would be capable of sustaining superconductivity as aresult of an electron-pairing mechanism involving electron-excitoninteractions (Little, W A: Possibility of Synthesizing an OrganicSuperconductor. Phys. Rev. 134(6A):A1416-A1424, 1964.). Little estimatedthat such a polymer, consisting of a conducting conjugated hydrocarbonbackbone and side chains in the form of highly polarizable dyemolecules, would be superconducting up to temperatures of the order2200° K. Such high temperatures would obviously not be realistic fororganic systems for reasons of thermal stability, but this estimate ofthe critical temperature does serve to indicate that the concept of theexistence of superconducting biopolymers at physiological temperatureslies well within the limit of the applicability of Little's theory. Theexistence of superconductivity in aromatic compounds was firstspeculated upon by London (London, F J: J. Phys. Radium 8:397, 1937);and Ladik et al. (Ladik, J; Biczo, G; Redly, J: Possibility ofSuperconductive-Type Enhanced Conductivity in DNA at Room Temperature.Phys. Rev. 188(2):710-715, 1969) have provided a theoretical basis forthe superconductive behavior of DNA.

Experimental evidence for high temperature superconduction in biologicalmolecules has been reported in several laboratories. Superconductivitywas deduced to occur in small domains included in the insulating bulk ofbile cholate test samples and so as to distinguish the effects from thatnormally found for the elemental superconductors, the cholates weredesignated a fractional or Type III superconductor. When small amountsof water are introduced into such materials the hydrophobic groups willtend to cluster together, and on subsequent slow desiccation smallmicelles will be formed. Such micelles are considered by Halpern andWolf to form superconducting domains.

Following the suggestion that enzymes and other biological materialspossess a metastable state with high dipole moment, Ahmed et al.investigated the dielectric and magnetic susceptibility properties ofthe dilute solutions of lysozyme (Ahmed, N A G; Calderwood, J H;Frohlich, H; Smith, C W: Evidence For Collective Magnetic Effects In AnEnzyme: Likelihood Of Room Temperature Superconductive Regions. Phys.Lett. 53A(2):129-130, 1975). It was found that magnetic fields of theorder of 0.6 tesla could produce very large changes (˜30%) in therelative permittivity of the solutions. This was suggestive ofsuperconductive behavior. It was suggested that in each lysozymemolecule there existed a small superconductive region with lineardimensions smaller than the London penetration depth, and that thecollective, superconductor-like, phenomena resulted from the formationof clusters of these small regions. This is similar to the cluster modelproposed for bile cholates. It was also suggested that not only thelysozyme molecules, but also water and ions may have played a role inthe establishment of the superconducting regions.

Other indirect evidence to suggest a biological role forsuperconductivity has been suggested by Cope (Cope, F W: Physiol. Chem.Phys. 3:403, 1971. Cope, F W: Physiol. Chem. Phys. 5:173, 1973) thathigh temperature superconduction may be expected in a sandwichconsisting of a thin conductive film or filament adjacent to adielectric layer. Cope considers that such superconducting sandwichesmay be ubiquitous in biological systems in the form of thin layers ofprotein and unsaturated lipids and hydrocarbon ring structures(conducting layer) adjacent to layers of water (polarizable dielectriclayer). Examples of such biological processes are impulse conductionvelocity in frog sciatic nerves and junctional electrical resistance ofcrayfish nerve. Such an effect can be well described in terms of a modelwhere the rate-limited biological process involves a superconductingtunneling current of single electrons and/or electron pairs (theJosephson current). It was suggested that as there was an apparentassociation of superconduction with growth, then the superconductivemicro-regions may have been individual purine and pyrimidine rings ofDNA and RNA with electron tunneling between rings along the length ofthe polymer chain. It was further suggested that superconductiveJosephson junctions in living systems may provide a physical mechanismwith more than enough sensitivity to explain how many biologicalorganisms are able to respond to weak magnetic fields.

Two-component plasmas (or more generally multi-component plasmas) as inan electron-hole liquid can support, other than the usual plasmon mode,a new collective mode called the “acoustic-plasmon mode”. Quantummechanical treatment of acoustic plasmons in one dimensional systemssuch as a long DNA molecule have attracted attention. Tanatar (Tanatar,B: Collective Modes in a Quasi-One Dimensional, Two-Component ElectronLiquid. Solid State Communications 92(8):699-702, 1994) stated that amotivation to study the acoustic plasmons in quasi-one-dimensionalelectron-hole systems comes from the fact that they may provide apairing mechanism like the BCS theory which leads to a superconductingtransition (Bardeen, J; Cooper, L N; Schrieffer, J R: Microscopic Theoryof Superconductivity. Phys. Rev. 106:162-164, 1957). Such an acousticplasmon mediated superconductivity has been proposed and elaborated fortwo-dimensional electron-hole liquids. Possibility of superconductivitydue to ordinary plasmons in quantum wires were also considered.Experiments to observe the acoustic plasmons in quasi-one-dimensionalstructures such as DNA, and their possible pairing mechanism leading tosuperconductivity would be most interesting.

In a preferred embodiment, uniaxially oriented liquid crystal conductingbiopolymers (protein and DNA) deposited by controlled EMOLE fabricationtechniques are used to produce a functional device. Such devices arethought to function via one or more superconducting mechanism(s)described above. For example, a GOD-DNA device generates an electronpair for each D-(+)-glucose molecule oxidized by the GOD protein enzymeheadgroup. Electron pair movement from protein FAD/FADH₂ prostheticgroup (redox center) through DNA quantum wire to underlyingsemiconductor substrate occurs via superconducting mechanism(s)described above. Many gated devices inject an electron pair, viasuperconducting mechanism, into p-type silicon of a p-n homojunctionsolar cell; combining with photogenerated majority carriers (holes), tolower the baseline photocurrent (I_(sc)). Decrease in photocurrent isdirectly proportional to D-(+)-glucose concentration. Change inphotocurrent occurs very rapidly and is accompanied by a near stepchange (see FIGS. 5 and 6 described below), resulting from differentialdevice injection of mobile charge carriers (electrons or holes) intop-type or n-type semiconductor substrate surfaces.

VII. APPLICATIONS

The sensors of this invention may be employed for a myriad ofapplications. For example, sensor based home health monitors will besimple-to-use, non-invasive and relatively inexpensive for use inmonitoring health conditions at home. Many physical functions—liverfunctions, ovulation, pregnancy, yeast infections, viral infections,bacterial infections, levels of cholesterol, triglycerides, sugar,hormones, drugs, water, salt, pH, sodium, and potassium—may be monitoredas easily as weight is now tracked by bathroom scales. The graying ofour population and the increasing costs of medical care will make theseproducts extremely popular.

In another preferred embodiment of this invention, a sensor, in aportable pen-based device, may be used to monitor compounds found in thehuman breath. The normal human breath contains hundreds of volatileorganic compounds that are reflective of the metabolic state of theperson. These volatile organic compounds have been quantitated by gaschromatographic (GC) and mass spectrometry (MS) methods in numerousstudies. Preferably, a sensor of this invention is exposed to exhaledbreath. In a preferred embodiment, the molecular recognition surface ofthe sensor will be alcohol dehydrogenase (ADH) which specifically bindsto ethanol; reduced mercaptoethanol, glutathione, or dithiothreitolwhich specifically binds to sulfur containing compounds; or a variety ofother molecular recognition sites to detect breath compounds readilyrecognized by those of skill in the art. The ADH-sensor will providepolice and highway patrol officers with a portable pen-basedbreathalyzer to validate drunk driving violations in the field. TheThio-sensor will provide individuals with a portable pen-basedbreathalyzer for discrete detection of halitosis (i.e., bad breath).

In another preferred embodiment, a device-based molecular recognitionchip (MRC) may be embedded in a magnetoosmotic (MOP) or electroosmoticpatch (EOP) which may be applied to the skin for real-time non-invasivequantitation of analytes found below the skin (i.e., analytes in bloodand deep anatomic structures). This is a non-invasive approach toanalyte quantitation alternate to exposure of the powered chip toinvasively drawn blood or other fluids described above. The MRC-MOP orMRC-EOP is suitable for non-invasive detection of small charged,uncharged, and zwitter ionic molecules and salts (i.e., analytes) lessthan 30,000 daltons found on the other side of complex synthetic orbiological barriers such as skin, adipose tissue, vascular walls (i.e.,venous and arterial vessel walls), isoparenteral walls, extravascularwalls, extracellular walls, cerebral vascular walls, blood brain barrier(BBB), and a variety of other man-made and natural membranes. The MOP,applies a combination of localized magnetic field gradients andhypertonic junctions to surfaces such as skin that it contacts. The EOP,applies a combination of localized electric field gradients andhypertonic junctions to surfaces such as skin that it contacts. Thispermits the MOP or EOP to draw analytes through semi-permeable membranesand skin for detection by the embedded MRC as described above.Preferably, the MRC-MOP or MRC-EOP may be equipped with a number ofmolecular recognition sites to perform a complete blood gas, bloodelectrolyte, hematocrit, blood sugar, and blood metabolite analysisnon-invasively (i.e., without drawing blood).

In a preferred embodiment, applied a.c or d.c. electric or magneticfields are utilized to change the orientational and positional order ofliquid crystal biological structures such as cellular membranes,cellular pores, blood vessels, skin, sweat glands, etc. to permitleakage of contained body analytes. A hypertonic junction will pull out,by means of a low chemical potential well, and concentrate leakyanalytes. The hypertonic junction is composed of a suitablepolyelectrolyte gel or solid polymer electrolyte (Gray, F M: SolidPolymer Electrolytes. Fundamentals and Technological Applications. VCHPublishers, Inc.; New York, Weinheim, Cambridge; 1991. Hara, M (ed.):Polyelectrolytes. Science and Technology. Marcel Dekker, Inc.; New York,Basel, Hong Kong; 1993) containing an embedded device-based molecularrecognition chip (MRC) for detection of specific analyte(s).

VIII. SCREENING AND ASSAYS

A semiconductor surface prepared according to the methods describedabove can be used to screen for ligands (i.e., analytes) having highaffinity for immobilized molecular recognition sites. A solutioncontaining an unmarked (not labeled) ligand is introduced to thesurface. Generally, little or no incubation time is required because ofthe immediate response of the molecular recognition chip (MRC) on theorder of milliseconds.

In a preferred embodiment, a semiconductor substrate prepared asdiscussed above is exposed to light while connected to a digitalmultimeter (DMM) which measures the short circuit volt/amp output (i.e.,V_(sc), I_(sc)) of the p-n junction solar cell substrate (as shown inFIG. 3). The powered chip can now be exposed to a solution containing anunmarked ligand. The unmarked ligand binds with high affinity to animmobilized molecular recognition site previously localized on the chipsurface. A square wave signal is generated by the powered chip in lessthan a few milliseconds in response to the binding event of the ligand(i.e., digital output from the chip). In a preferred embodiment,D-(+)-glucose is applied to the surface of a glucose oxidase-molecularrecognition chip (GOD-Chip). The light powered GOD-Chip produces asquare wave response in volt/amp output proportional to the appliedD-(+)-glucose concentration (see FIGS. 5 and 6 described below). Thisreflects a change in the short circuit output of the p-n junction solarcell substrate due to electron tunneling from the molecular recognitionsurface (i.e., GOD) through the highly conductive polymer monolayer(i.e., liquid-crystal oriented B-DNA) to the p-type surface of the p-njunction solar cell. Carrier injection of electrons by a device into thep-type layer of a powered solar cell substrate interrupts the baselineshort circuit photovoltage and photocurrent (i.e., V_(sc), I_(sc)) ofthis simple p-n junction, rectifying diode device. The amount ofelectrons injected into the p-type surface of the powered chip isproportional to the amount of D-(+)-glucose binding to the GOD, which isreflected in a proportional digital square wave output by the GOD-Chip(FIG. 5 and FIG. 6). Device injection of electrons directly into p-typesilicon eliminates or lowers the photocurrent by combination withphotogenerated carrier holes before than can recombine, via shortcircuit wire, with photogenerated carrier electrons from the n-typelayer. Concomitantly with lowered photocurrent resulting from carrierhole removal from the p-type layer, the continued build-up ofphotogenerated carrier electrons in the n-type layer increases measuredphotovoltage of the circuit (FIG. 6).

In this embodiment, a simple digital multimeter (volt/amp) is employedto measure the digital output of the GOD-Chip. Therefore, single andmultiple IC arrays described above may be configured in pen-baseddigital meters, hand-held digital meters, clinical lab-basedinstruments, digital wireless implantable medical devices, andindustrial-based digital devices which measure real-time molecularbinding events and constants of analytes. A simple calibration curve foreach chip can be used to determine the concentration of unknown samples.Calibrated chips are not affected by altitude, humidity, O₂ partialpressure, diffusional electron acceptor mediators, or application of thesample. These problems of the prior art, have been overcome in thepresent invention because electron transfer rates of the molecular wireinterconnects are orders of magnitude greater than enzymatic reactionrates, and electron transfer rates of diffusional redox mediators suchas O₂ and other small molecule inorganic, organometallic, and organiccompounds used in amperometric detection methods. A forward electrontransfer rate constant (k_(f)>10⁷ s⁻¹ Å⁻²) may be very high because ofthe quantum-wire nature (i.e., defined electronic energy levels) of theconductive polymer interconnects. Connecting polymers may also bereversibly switched between conductive and insulative states byoxidation or reduction.

IX. EXAMPLES

The following examples of preferred embodiments of the present inventionare presented by way of illustration only and do not suggest that theabove-described methods and compositions are in any way limited by thespecific examples set forth below.

Example A Preparation of Polycrystalline Silicon p-n Junction Solar Cell

A commercial polycrystalline silicon p-n junction solar cell chip (a0.1799 g and 1.75 cm²) from Edmund Scientific, Barrington, N.J.08007-1380 (Stock Nos. 35,220 and 35,221) was exposed to 980 Lux lightintensity from a F15T8/CW Westinghouse bulb at 25° C. With the dark blueemitter surface facing the light source, the short circuit DC output ofthe dry solar cell substrate was measured by a digital multimeter (DMM)(Extech Instruments; Model No. 383273). Measured DC output was 121 mVand 98 uA. The solar cell chip was then washed with analytical reagentelectronic-grade solvents: i) acetone; ii) methanol; iii) 18 Mohm H₂O;and iv) methanol. It was allowed to dry in a dust-free environment. Thisprepared the semiconductor p-n type surfaces for electroplating.

Example B Electrodeposition of DNA onto a Polycrystalline Silicon p-nJunction Solar Cell

A DNA electroplating solution was prepared using 18 Mohm sterile wateras the solvent. 0.1062 g of DNA (degraded free acid from Herring sperm)was added to 100 ml of water. The pH of the resulting solution was˜2.00. The pH was adjusted to ˜7.00 with NaOH and HCl. No buffer wasadded to the electroplating solution. The final salt/electrolyteconcentration was <150 mM. 1.00 mL of methanol was added to the DNAelectroplating solution and mixed thoroughly. The dry solar cell chipfrom Example A produced a short circuit output of 152 mV and 136 uA whenexposed to 1700 Lux light intensity generated from two F15T8/CWWestinghouse bulbs at 25° C. The dry solar cell chip from Example A wassubmerged in the DNA electroplating bath at 25° C. with dark blueemitter surface exposed to 1700 Lux light intensity generated from twoF15T8/CW Westinghouse bulbs. After ˜5.50 hours, the solar cell chip wasremoved from the DNA bath and placed on a paper towel to dry. The darkblue emitter surface was exposed to the 1700 Lux light source during thedrying process which took ˜12.00 hours at 25° C. in air. DNAelectroplated on the back or silvery side of the solar cell chip (i.e.,p-type silicon) as evidenced by a white coating visible to the eye. Onthe dark blue emitter surface (i.e., n-type silicon) no significantcoating was observed. The pH of the DNA electroplating bath remained˜7.00 after the electroplating process was complete. The electrochemicalpotential at the plating surface was about 150 mV and the currentdensity was about 77 μA cm⁻².

Example C Electrodeposition of Glucose Oxidase (GOD) onto a DNA-CoatedPolycrystalline Silicon p-n Junction Solar Cell

A glucose oxidase (GOD) electroplating solution was prepared using 18Mohm sterile water as the solvent. 0.0092 g of glucose oxidase (EC1.1.3.4; ˜1,000 units) was added to 100 ml of water. The pH of theresulting solution was ˜6.00. No buffer or further adjustment of pH wasnecessary. 1.00 ml of methanol was added to the GOD electroplatingsolution and mixed thoroughly. Next, the DNA-coated polycrystallinesilicon p-n junction solar cell from Example B was submerged in the GODelectroplating bath with dark blue emitter surface exposed to 1700 Luxlight intensity from two F15T8/CW Westinghouse bulbs at 25° C. for ˜8.10hours. The solar cell chip was removed from the GOD bath and placed on apaper towel to dry. The dark blue emitter surface was exposed to the1700 Lux light source during the drying process which took ˜12.00 hoursat 25° C. in air. GOD electroplated on the back or silvery side of thesolar cell chip (i.e., p-type silicon) as evidenced by a yellow-orangeprecipitate visible to the eye. The yellow-orange GOD precipitate was inthe same area of the chip overlapping the white DNA precipitate fromexample B. On the dark blue emitter surface (i.e., n-type silicon) nosignificant coating was observed. The pH of the GOD electroplating bathremained ˜6.00 after the electroplating process was complete. TheGOD-DNA-Chip was removed from the light and put under parafilm toprotect and store until use. The electrochemical potential at theplating surface was about 150 mV and the current density was about 77 μAcm⁻².

Example D Detection of D-(+)-glucose on a GOD-DNA-Chip

Coating/electroplating of the solar cell chip from example A did notchange the electronic output characteristics of the device prior totesting with the D-(+)-glucose ligand.

The dry GOD-DNA-Chip from example C was placed with the silver GOD-DNAcoated surface (i.e., p-type silicon) facing a F15T8/CW Westinghousebulb. A red (positive) test lead of a digital multimeter (DMM) ExtechInstruments; Model No. 383273) was connected to the dark blue emitter(i.e., n-type) surface and the black (negative) test lead was connectedto the p-type GOD-DNA coated surface facing the light (FIG. 3). Theintensity of light was adjusted to produce a baseline short circuitcurrent of approximately −60 uA (FIG. 5). After several minutes, a drop(˜0.100 mL) of a sterile D-(+)-glucose standard (63 mg/dL) was placed onthe powered GOD-DNA-Chip resulting in a large square wave amplitudechange of approximately +51 uA reaching a new baseline of approximately−8 uA (FIG. 5). This is consistent with approximately 2.00×10¹⁷ glucosemolecules being applied to the chip in a 1 cm² area generating themaximum current expected from a monolayer of well connected GOD. Glucoseoxidase (GOD) turns over at ambient temperature at a rate of ˜10² s⁻¹,i.e., it produces about 200 transferable electrons/s. Because it has aradius of ˜43 Å, there can be up to 1.7×10¹² enzyme molecules on theelectrode surface. The current density, when all redox centers areelectrically well connected to the electrode, may thus reach about3.4×10¹⁴ electrons s⁻¹ cm⁻², or 53 uA cm⁻² (Heller, A: Electrical Wiringof Redox Enzymes. Acc. Chem. Res. 23(5):128-134, 1990).

Another test of GOD-DNA-Chip performance at different D-(+)-glucoseconcentration levels is demonstrated in FIG. 6. “Level 1” and “level 2”are sterile D-(+)-glucose standards (˜63 and 20 mg/dL respectively). Adrop of “level 1” D-(+)-glucose standard produces the first square wave;followed by washing with H₂O and application of the lower “level 2”D-(+)-glucose concentration. Square wave amplitude responses aredirectly proportional to the D-glucose concentrations applied to thechip. Washing the GOD-DNA-Chip of ligand D-(+)-glucose with H₂O returnsthe chip to its baseline voltage/current (FIG. 5 and FIG. 6).

Example E Electrodeposition of Glucose Dehydrogenase (GDH) onto aDNA-Coated Polycrystalline Silicon p-n Junction Solar Cell

A DNA-coated polycrystalline silicon solar cell was prepared in a mannersimilar to that explained above in Examples A and B. The differenceswere as follows:

1. A commercial polycrystalline silicon solar cell chip 0.0280 g and0.4059 cm² was used as the semiconductor substrate.

2. The dry solar cell chip from 1 (above) produced a short circuitoutput of 43.55 mV and 36.35 microAmperes when exposed to 1700 Lux lightintensity generated from two F15T8/CW Westinghouse fluorescent bulbs at25 degrees Centigrade.

3. The dry solar cell chip was submerged into 300 microLiters of theDNA/EMOLE™ electroplating bath at 25 degrees Centigrade with the darkblue emitter surface of the chip exposed to 1700 Lux light intensitygenerated by two F15T/CW Westinghouse fluorescent bulbs.

4. After 38.75 hours, the solar cell chip was removed from theDNA/EMOLE™ electroplating bath.

5. The DNA-chip was dried under 1700 Lux light intensity generated bytwo F15T8/CW Westinghouse fluorescent bulbs and an umbrella of N₂ gasfor 2.50 hours at 25 degrees Centigrade.

6. The electrochemical potential at the plating surface of the siliconsemiconductor substrate was about 44 mV and the current density wasabout 89 microAmperes/cm⁻².

A glucose dehydrogenase (GDH) electroplating solution was prepared using18 Mohm sterile water as the solvent. 0.0046 g of glucose dehydrogenase(EC 1.1.1.119; 50 units) was added to 7.50 mL of water. The pH of theresulting solution was 6.728. No addition of buffer or furtheradjustment of pH was necessary. 75 microLiters of methanol was added tothe GDH electroplating solution and mixed thoroughly. Next, the dryDNA-chip from above was submerged into 300 microLiters of the GDH/EMOLE™electroplating bath with the dark blue emitter surface of the chipexposed to the 1700 Lux light intensity generated by two F15T8/CWWestinghouse fluorescent bulbs at 25 degrees Centigrade for 26.75 hours.The solar cell chip was removed from the GDH/EMOLE™ electroplating bath.The GDH-DNA-chip was dried under 1700 Lux light intensity generated bytwo F15T8/CW Westinghouse fluorescent bulbs and an umbrella of N₂ gasfor 5.00 hours at 25 degrees Centigrade. The GDH-DNA-chip was removedfrom the light and put in a desiccator box to protect and store untiluse. The electrochemical potential at the plating surface of the siliconsemiconductor substrate was about 44 mV and the current density wasabout 89 microAmperes/cm⁻².

Example F Detection of D-(+)-Glucose on a GDH-DNA-Chip

As in the GOD examples, EMOLE™ coating/electroplating of the solar cellchip did not change the electronic output characteristics of the deviceprior to testing with the D-(+)-glucose ligand.

The dry GDH-DNA-chip from Example E was placed with the silver GDH-DNAcoated surface (i.e., p-type silicon) facing a F15T8/CW Westinghousefluorescent bulb. A black (negative) test lead of a digital multimeter(Hewlett-Packard Model 34970A) was connected to the dark blue emitter(i.e., n-type silicon) surface and the red (positive) test lead wasconnected to the p-type GDH-DNA coated surface facing the light. Theintensity of the light was adjusted to produce a baseline short circuitcurrent of approximately +65 microAmperes (lower curve of FIG. 7) andbaseline potential of approximately +78 mV (upper curve of FIG. 7).After a few minutes, 5 microLiters of sterile D-(+)-glucose (60 mg/dL)in saline sodium phosphate buffer (1×SSP, pH 7.323) was dropped on theGDH-DNA-chip resulting in an immediate large square wave amplitudechange of approximately +6.5 microAmperes and +7.5 mV reaching newbaselines of approximately +71 microAmperes and +85 mV, respectively(FIG. 7).

The above examples employing GOD and GDH serve to illustrate the utilityand wide applicability of the present invention. While both GOD (EC1.1.3.4) and GDH (1.1.1.1.19) oxidize D-(+)-glucose to D-gluconolactone,they are very different enzymes.

GOD (EC 1.1.3.4) is widespread among fungi. GOD is a FAD containingflavoprotein and glycoprotein with a molecular mass of 160,000 daltons.GOD contains two moles of FAD cofactor per mole of enzyme and 16%carbohydrate, the carbohydrate chains are not directly involved incatalysis. The specificity of GOD is very high, the beta-form of glucoseis oxidized 157 times more rapidly than the alpha-form and of othersubstrates examined only 2-deoxy-D-glucose and 6-deoxy-D-glucose wereoxidized at rates greater than 10% of that of D-glucose. O₂ is thenatural electron acceptor of this enzyme producing H₂O₂.

The NAD(P)-dependent GDH (EC 1.1.1.119) occurs in photoautotrophicprokaryotes such as strains of bacteria capable of forming on glucose inthe dark. In addition to oxidizing D-(+)-glucose (alpha- andbeta-forms), NAD(P)-dependent GDH also oxidizes D-mannose,2-deoxy-D-glucose, and 2-amino-2-deoxy-D-mannose. NAD(P)-dependent GDHis not a flavoprotein or glycoprotein and has unusual specificity; itdoes not oxidize aldopentoses and is completely inactive with NAD⁺ or O₂as electron acceptors. Instead it very specifically requires NAD(P)⁺ asits electron acceptor, producing NAD(P)H+H⁺. The molecular mass of theenzyme is approximately 230,000 daltons. Oxidation of D-mannose is arelatively unusual feature of aldose dehydrogenases obtained fromvarious biological sources.

If NAD(P)⁺ is not available in solution, GDH (EC 1.1.1.119) will notoxidize D-(+)-glucose. NAD(P)⁺ was not added to the GDH-DNA-chip testsolution in Example F which nevertheless rapidly oxidized addedD-(+)-glucose indicating that the DNA molecular wire of this device hasreplaced diffusible NAD(P)⁺, not present in the test solution, as a“hard wired” conduit for direct electron transfer from the attachedcatalytic headgroup GDH enzyme to the silicon semiconductor substrate.

As mentioned above, GOD in its native state oxidizes D-glucose throughits FAD/FADH₂ redox center. This involves two electrons and two hydrogenions being transferred to the FAD prosthetic group which is tightlybound to the enzyme. Normally, in the absence of a sensor mediator, theGOD-FADH₂ complex is reoxidized by atmospheric oxygen (i.e., O₂) toGOD-FAD complex to complete the catalytic reaction cycle. GDH (EC1.1.1.119), by contrast, is not a FAD containing flavoprotein. GDH (EC1.1.1.119) oxidizes D-glucose through a different redox center utilizingdiffusible NAD(P)⁺ coenzyme in stoichiometric amounts which is broughtinto play during the catalytic mechanism of oxidation and electrontransfer producing D-gluconolactone and NAD(P)H+H⁺. The reduced coenzymeNAD(P)H is not recycled nor reoxidized by molecular oxygen (i.e., O₂)(as with GOD-FADH₂) so that enough expensive NAD(P)⁺ coenzyme must beadded in the beginning to drive the biocatalytic oxidation of glucose.Thus, glucose sensors relying on GDH (EC 1.1.1.119) are not sensitive tooxygen partial pressure, unlike GOD-based glucose sensors. Stillfurther, the GOD and GDH amino acid sequences are completely different.The GDH enzyme has a molecular mass of approximately 230,000 daltonswhile the GOD enzyme has a molecular mass of approximately 160,000daltons. Thus, the above examples demonstrate that the invention can beapplied to widely different molecular recognition headgroups.

CONCLUSION

Various references have been cited in this specification. Each of thesereferences is incorporated herein by reference for all purposes.

The invention has been described primarily with reference to the use ofelectrochemical deposition of liquid-crystal conductive polymers andmolecular recognition surfaces, but it will be readily recognized bythose of skill in the art that other types of deposition, conductivewiring, and substrates can be used. Various forms of patternedelectrochemical and chemical deposition may be used. Many types of p-nhetero- or homojunction semiconductor substrates may be used. Thesubstrate may be powered by broad spectrum light, light emitting diodes(LED), lasers, solar radiation, uv radiation, vis radiation, infraredradiation, x-rays, gamma rays, radioactivity, thermally, or by anyexternal supplied nuclear or electromagnetic energy greater than thesubstrate bandgap to provide patterned areas of electrochemicaldeposition and to power the completed device.

It is understood that the above description is intended to beillustrative and not restrictive. Many embodiments will be apparent tothose of skill in the art upon reviewing the above description. Thescope of the invention should, therefore, be determined not withreference to the above description, but should instead be determinedwith reference to the appended claims, along with the full scope ofequivalents to which such claims are entitled.

1. A sensor for sensing the presence of an analyte component, the sensorcomprising: a plurality of conductive polymer strands each having atleast a first end and a second end, and being in contact with amolecular insulator that isolates one or more of said conductive polymerstands from one or more other conductive polymer strands on the sensor;a plurality of molecular recognition headgroups having an affinity forsaid analyte component and being attached to said conductive polymerstrands such that when the analyte interacts with the molecularrecognition headgroup one or more mobile charge carriers are transferredto a conductive polymer strand attached to said headgroup; and anelectrode substrate attached to said conductive polymer strands at saidsecond ends configured to report to an electronic circuit reception ofmobile charge carriers from said conductive polymer strands whereby thepresence of said analyte component is sensed.
 2. The sensor of claim 1,wherein the conductive polymer strands are aligned in a substantiallynon-random orientation.
 3. The sensor of claim 1, wherein the conductivepolymer strands are aligned in a substantially uniaxial orientation. 4.The sensor of claim 1, wherein the molecular recognition headgroupsparticipate in a redox reaction when contacting a molecule of saidanalyte component, and wherein when said redox reaction occurs at aheadgroup, a mobile charge carrier is transferred directly to aconductive polymer strand attached to said headgroup, without redoxreaction in the polymer strand.
 5. The sensor of claim 1, wherein theplurality of conductive polymer strands are multi-stranded nucleic acidstrands.
 6. The sensor of claim 1, wherein the plurality of conductivepolymer strands have an orientation that is preferentially orthogonal tothe electrode substrate.
 7. The sensor of claim 1, wherein the pluralityof molecular recognition headgroups are selected from the groupconsisting of oxidoreductases, immunoglobulins and catalytic antibodies.8. The sensor of claim 1, wherein the plurality of molecular recognitionheadgroups are chemically homogeneous.
 9. The sensor of claim 1, whereinthe plurality of molecular recognition headgroups are chemicallyinhomogeneous.
 10. The sensor of claim 9, wherein the sensor includes afirst region on said electrode substrate where a first group ofchemically homogeneous molecular recognition headgroups is located andsecond region on said electrode substrate where a second group ofchemically homogeneous molecular recognition headgroups is located, andwherein the first and second regions are separately addressable.
 11. Thesensor of claim 10, wherein conductive polymer strands of the firstregion are electrically isolated from conductive polymer strands of thesecond regions by said molecular insulator.
 12. The sensor of claim 1,wherein the molecular recognition headgroups include at least one ofglucose oxidase and glucose dehydrogenase.
 13. A method of forming asensor capable of sensing the presence of an analyte component, themethod comprising: affixing conductive polymer strands or precursors ofsaid conductive polymer strands to a sensor substrate; affixingmolecular recognition headgroups to said affixed conductive polymerstrands, whereby a sensor structure is formed having said substrateaffixed to said conductive polymer strands and said molecularrecognition headgroups also affixed to said conductive polymer strands;and contacting at least the conductive polymer strands with a molecularinsulator that isolates one or more of said conductive polymer standsfrom one or more other conductive polymer strands on the sensor.
 14. Themethod of claim 13, wherein affixing the conductive polymer strands orprecursors of said conductive polymer strands to a sensor substratecomprises contacting a sensor substrate with a first medium containingconductive polymer strands or precursors of said conductive polymerstrands.
 15. The method of claim 14, wherein the step of applying apotential is performed at a potential which causes said affixedconductive polymer strands to be oriented in a non-random direction. 16.The method of claim 13, wherein affixing the molecular recognitionheadgroups to said affixed conductive polymer strands comprises applyinga potential to said substrate sufficient to affix said molecularrecognition headgroups to said affixed conductive polymer strands. 17.The method of claim 13, wherein the conductive polymer strands aremulti-stranded nucleic acids.
 18. The method of claim 13, wherein thesensor substrate is a device element of a device on semiconductor chip.19. The method of claim 13, further comprising isolating a region of thesensor substrate prior to affixing the molecular recognition headgroupsto said affixed conductive polymer strands, such that the molecularrecognition headgroups are deposited only in the isolated region.
 20. Amethod of sensing the presence of an analyte component in an analytewith a sensor including (i) a plurality of conductive polymer strandseach having at least a first end and a second end, and being in contactwith a molecular insulator that isolates one or more of said conductivepolymer stands from one or more other conductive polymer strands on thesensor, (ii) a plurality of molecular recognition headgroups having anaffinity for said analyte component and being attached to saidconductive polymer strands such that when the analyte interacts with themolecular recognition headgroup one or more mobile charge carriers aretransferred to a conductive polymer strand attached to said headgroup,and (iii) an electrode substrate attached to said conductive polymerstrands at said second ends, the method comprising: contacting saidmolecular recognition headgroups with said analyte; and determiningwhether mobile charge carriers have been transferred to said electrodesubstrate resulting from mobile charge carriers transferred by saidconductive polymer strands, at least some of which are isolated from oneanother by said molecular insulator, to said electrode substrate tothereby sense the presence of the analyte component.